Systems and methods for controlling an implantable blood pump

ABSTRACT

Systems and methods for controlling an implantable pump are provided. For example, the exemplary controller for controlling the implantable pump may only rely on the actuator&#39;s current measurement. The controller is robust to pressure and flow changes inside the pump head, and allows fast change of pump&#39;s operation point. For example, the controller includes, a two stage, nonlinear position observer module based on a reduced order model of the electromagnetic actuator. The controller includes an algorithm that estimates the position of the moving component of the implantable pump based on the actuator&#39;s current measurement and adjusts operation of the pump accordingly. Alternatively, the controller may rely on position measurements and/or velocity estimations.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of International PCT PatentApplication Serial No. PCT/IB2020/052337, filed Mar. 13, 2020, whichclaims the benefit of priority of U.S. Provisional Patent ApplicationNo. 62/819,436, filed Mar. 15, 2019, the entire contents of each ofwhich are incorporated herein by reference. This application is alsorelated to U.S. patent application Ser. No. 16/819,021, filed Mar. 13,2020, now U.S. Pat. No. 10,799,625, which claims the benefit of priorityof U.S. Provisional Patent Application No. 62/819,436, filed Mar. 15,2019, the entire contents of each of which are incorporated herein byreference.

FIELD OF THE INVENTION

The present invention relates generally to implantable heart pumpshaving an undulating membrane with improved hydraulic performancedesigned to reduce hemolysis and platelet activation and moreparticularly to controlling the implantable pump.

BACKGROUND

The human heart is comprised of four major chambers with two ventriclesand two atria. Generally, the right-side heart receives oxygen-poorblood from the body into the right atrium and pumps it via the rightventricle to the lungs. The left-side heart receives oxygen-rich bloodfrom the lungs into the left atrium and pumps it via the left ventricleto the aorta for distribution throughout the body. Due to any of anumber of illnesses, including coronary artery disease, high bloodpressure (hypertension), valvular regurgitation and calcification,damage to the heart muscle as a result of infarction or ischemia,myocarditis, congenital heart defects, abnormal heart rhythms or variousinfectious diseases, the left ventricle may be rendered less effectiveand thus unable to pump oxygenated blood throughout the body.

The Centers for Disease Control and Prevention (CDC) estimate that about5.1 million people in the United States suffer from some form of heartfailure. Heart failure is generally categorized into four differentstages with the most severe being end stage heart failure. End stageheart failure may be diagnosed where a patient has heart failuresymptoms at rest in spite of medical treatment. Patients at this stagemay have systolic heart failure, characterized by decreasing ejectionfraction. In patients with systolic heart failure, the walls of theventricle, which are typically thick in a healthy patient, become thinand weak. Consequently, during systole a reduced volume of oxygenatedblood is ejected into circulation, a situation that continues in adownward spiral until death. A patient diagnosed with end stage heartfailure has a one-year mortality rate of approximately 50%.

For patients that have reached end stage heart failure, treatmentoptions are limited. In addition to continued use of drug therapycommonly prescribed during earlier stages of heart failure, the typicalrecommend is cardiac transplantation and implantation of a mechanicalassist device. While a cardiac transplant may significantly prolong thepatient's life beyond the one year mortality rate, patients frequentlyexpire while on a waitlist for months and sometimes years awaiting asuitable donor heart. Presently, the only alternative to a cardiactransplant is a mechanical implant. While in recent years mechanicalimplants have improved in design, typically such implants will prolong apatient's life by a few years at most, and include a number ofco-morbidities.

One type of mechanical implant often used for patients with end stageheart failure is a left ventricular assist device (LVAD). The LVAD is asurgically implanted pump that draws oxygenated blood from the leftventricle and pumps it directly to the aorta, thereby off-loading(reducing) the pumping work of the left ventricle. LVADs typically areused either as “bridge-to-transplant therapy” or “destination therapy.”When used for bridge-to-transplant therapy, the LVAD is used to prolongthe life of a patient who is waiting for a heart transplant. When apatient is not suitable for a heart transplant, the LVAD may be used asa destination therapy to prolong the life, or improve the quality oflife, of the patient, but generally such prolongation is for only acouple years.

Generally, a LVAD includes an inlet cannula, a pump, and an outletcannula, and is coupled to an extracorporeal battery and control unit.The inlet cannula typically directly connected to the left ventricle,e.g., at the apex, and delivers blood from the left ventricle to thepump. The outlet cannula typically connected to the aorta distal to theaortic valve, delivers blood from the pump to the aorta. Typically, theoutlet cannula of the pump is extended using a hose-type structure, suchas a Dacron graft, to reach a proper delivery location on the aorta.Early LVAD designs were of the reciprocating type but more recentlyrotary and centrifugal pumps have been used.

U.S. Pat. No. 4,277,706 to Isaacson, entitled “Actuator for Heart Pump,”describes a LVAD having a reciprocating pump. The pump described in theIsaacson patent includes a housing having an inlet and an outlet, acavity in the interior of the pump connected to the inlet and theoutlet, a flexible diaphragm that extends across the cavity, a platesecured to the diaphragm, and a ball screw that is configured to bereciprocated to drive the plate and connected diaphragm from one end ofthe cavity to the other end to simulate systole and diastole. The ballscrew is actuated by a direct current motor. The Isaacson patent alsodescribes a controller configured to manage the revolutions of the ballscrew to control the starting, stopping and reversal of directions tocontrol blood flow in and out of the pump.

Previously-known reciprocating pump LVADs have a number of drawbacks.Such pumps often are bulky, heavy and may require removal of bones andtissue in the chest for implantation. They also require a significantamount of energy to displace the blood by compressing the cavity.Moreover, the pump subjects the blood to significant pressurefluctuations as it passes through the pump, resulting in high shearforces and risk of hemolysis. These pressure fluctuations may beexaggerated at higher blood flow rates. Further, depending on thegeometry of the pump, areas of little or no flow may result in flowstagnation, which can lead to thrombus formation and potentially fatalmedical conditions, such as stroke. Finally, the positive displacementpumps like the one described in the Isaacson patent are incapable ofachieving pulsatility similar to that of the natural heart, e.g.,roughly 60 to 100 beats per minute, while maintaining physiologicalpressure gradients.

LVADs utilizing rotary and centrifugal configurations also are known.For example, U.S. Pat. No. 3,608,088 to Reich, entitled “ImplantableBlood Pump,” describes a centrifugal pump to assist a failing heart. TheReich patent describes a centrifugal pump having an inlet connected to arigid cannula that is coupled to the left ventricular cavity and aDacron graft extending from the pump diffuser to the aorta. A pumpincludes an impeller that is rotated at high speeds to accelerate blood,and simulated pulsations of the natural heart by changing rotationspeeds or introducing a fluid oscillator.

U.S. Pat. No. 5,370,509 to Golding, entitled “Sealless Rotodynamic Pumpwith Fluid Bearing,” describes an axial blood pump capable for use as aheart pump. One embodiment described involves an axial flow blood pumpwith impeller blades that are aligned with the axes of the blood inletand blood outlet. U.S. Pat. No. 5,588,812 to Taylor, entitled“Implantable Electrical Axial-Flow Blood Pump,” describes an axial flowblood pump similar to that of the Golding patent. The pump described inthe Taylor patent has a pump housing that defines a cylindrical bloodconduit through which blood is pumped from the inlet to the outlet, androtor blades that rotate along the axis of the pump to accelerate bloodflowing through the blood conduit.

While previously-known LVAD devices have improved, those pump designsare not without problems. Like reciprocating pumps, rotary andcentrifugal pumps are often bulky and difficult to implant. Rotarypumps, while mechanically different from positive displacement pumps,also exhibit undesirable characteristics. Like positive displacementpumps, rotary pumps apply significant shear forces to the blood, therebyposing a risk of hemolysis and platelet activation. The very nature of adisk or blade rotating about an axis results in areas of high velocityand low velocity as well as vibration and heat generation. Specifically,the area near the edge of the disk or blade furthest from the axis ofrotation experiences higher angular velocity and thus flow rate than thearea closest to the axis of rotation. The resulting radial velocityprofile along the rotating blade results in high shear forces beingapplied to the blood. In addition, stagnation or low flow rates near theaxis of rotation may result in thrombus formation.

While centrifugal pumps may be capable generating pulsatile flow byvarying the speed of rotation of the associated disk or blades, thisonly exacerbates the problems resulting from steep radial velocityprofiles and high shear force. In common practice, the output ofcurrently available rotary pumps, measured as flow rate against a givenhead pressure, is controlled by changing the rotational speed of thepump. Given the mass of the rotating member, the angular velocity of therotating member, and the resulting inertia, a change in rotational speedcannot be instantaneous but instead must be gradual. Accordingly, whilecentrifugal pumps can mimic a pulsatile flow with gradual speed changes,the resulting pulse is not “on-demand” and does not resemble a typicalphysiological pulse.

Moreover, rotary pumps typically result in the application ofnon-physiologic pressures on the blood. Such high operating pressureshave the unwanted effect of overextending blood vessels, which in thepresence of continuous flow can cause the blood vessels to fibrose andbecome inelastic. This in turn can lead to loss of resilience in thecirculatory system, promoting calcification and plaque formation.Further, if the rotational speed of a pump is varied to simulatepulsatile flow or increase flow rate, the rotary pump is less likely tobe operated at its optimal operating point, reducing efficiency andincreasing energy losses and heat generation.

LVADs may also be configured to increase blood flow to match the demandof the patient. Numerous publications and patents describe methods foradjusting LVAD pump flow to match that demanded by the patient. Forexample, U.S. Pat. No. 7,520,850 to Brockway, entitled “Feedback controland ventricular assist devices,” describes systems and methods foremploying pressure feedback to control a ventricular assist device. Thesystem described in the Brockway patent attempts to maintain a constantfilling of the ventricle by measuring ventricular pressure and/orventricular volume. While such systems can achieve flow rates as high as8 or 9 liters per minute, these flow rates generally are outside of theefficient range of operation for current rotary pumps, which aretypically tuned to operate in a range of 4 to 6 liters per minute. Thus,increasing the flow rate in rotary pumps to match patient demandedresults in non-optimal pump performance.

Pumps other than of the rotary and positive displacement types are knownin the art for displacing fluid. For example, U.S. Pat. Nos. 6,361,284and 6,659,740, both to Drevet, entitled “Vibrating Membrane FluidCirculator,” describe pumps in which a deformable membrane is vibratedto propel fluid through a pump housing. In these patents, vibratorymotion applied to the deformable membrane causes wave-like undulationsin the membrane that propel the fluid along a channel. Different flowrates may be achieved by controlling the excitation applied to themembrane.

U.S. Pat. No. 7,323,961 to Drevet, entitled “Electromagnetic Machinewith a Deformable Membrane,” describes a device in which a membrane iscoupled in tension along its outer edge to an electromagnetic devicearranged to rotate around the membrane. As the electromagnetic devicerotates, the outer edge of the membrane is deflected slightly in adirection normal to the plane of the membrane. These deflections inducea wave-like undulation in the membrane that may be used to move a fluidin contact with the membrane.

U.S. Pat. No. 9,080,564 to Drevet, entitled “Diaphragm Circulator,”describes a tensioned deformable membrane in which undulations arecreated by electromechanically moving a magnetized ring, attached to anouter edge of a deformable membrane, over a coil. Axial displacement ofmagnetized ring causes undulations of membrane. Like in the '961 patent,the membrane undulations can be controlled by manipulating the magneticattraction. U.S. Pat. No. 8,714,944 to Drevet, entitled “Diaphragm pumpwith a Crinkle Diaphragm of Improved Efficiency” and U.S. Pat. No.8,834,136 to Drevet, entitled “Crinkle Diaphragm Pump” teach similartypes of vibrating membrane pumps.

None of the foregoing patents to Drevet describe a vibratory membranepump suitable for use in a biological setting, or capable of pumpingblood over extended periods that present a low risk of flow stagnationleading to thrombus formation.

U.S. Patent Publication Nos. 2017/0290966 and 2017/0290967 toBotterbusch, the entire contents of each of which are incorporatedherein by reference, describe implantable cardiovascular blood pumpshaving a flexible membrane coupled to an electromagnetic actuatorassembly that causes wavelike undulations to propagate along theflexible membrane to propel blood through the pump while avoidingthrombus formation, hemolysis and/or platelet activation. TheBotterbusch pumps generate hydraulic power—flow and pressure—bytranslating the linear motion of the electromagnetic actuator, to theflexible membrane, which deforms through its interaction with the blood,translating energy to the blood. The flexible membrane is oriented at a90° angle to the motion of the linear actuator such that the outer edgeof the membrane is the first element to engage the blood. As a result,there is a risk of energy loss at the inlet to the membrane, whichaffects the hydraulic power generation by the pump.

What is needed is an energy efficient implantable pump having lightweight, small size, and fast start and stop response that can operateefficiently and with improved hydraulic performance and minimal blooddamage over a wide range of flow rates.

The design of such an energy efficient implantable pump that fulfils allthe requirements mentioned above poses many challenges in terms ofmechanical design and manufacturing process. It is also a challenge froma control perspective because unlike rotary pumps, the operation pointof a vibrating membrane pump is set by the frequency and amplitude ofmembrane excitation. Indeed, the higher the frequency or the stroke ofthe undulation is, the higher the pressure head of the implantable pumpwill be. The stroke needs to be set accurately with sufficient speed tobe able to switch the operating point of the pump fast enough torecreate a sufficient pulse that is synchronized to heartbeats. At thesame time, the stroke must be restrained so as not to damage themembrane, blood, or the internal spring components of the pump byexcessive stress. This phenomenon can be caused by overpowering theactuator or by the effect of perturbation forces induced by theremaining activity of the left ventricle. Due to the specific medium(blood) in which the pump is operating, it may be preferred to avoidadding position, velocity, or acceleration sensors that wouldsignificantly increase the complexity and size of a pump that is alreadydifficult to design.

Attempts to bypass the use of motion sensors include those that measurecurrent ripple generated by a pulse-width modulation (PWM) voltage inputto estimate an equivalent circuit inductance that is related to themagnet position. (See, e.g., M. F. Rahman, et al., Position estimationin solenoid actuators, IEEE Transactions on Industry Applications, vol.32, n. 3, p. 552-559, June 1996). This method only works if the magneticparts' velocity is close to zero which is not the case of vibratingmembrane pump that operates at frequencies close to 100 Hz. Othersmethods compute the back electromotive force (back EMF proportional tovelocity) from an inverted equivalent electric circuit and directlyintegrate the estimated speed to get the position. (See, e.g., J. Zhang,et al., Study on Self-Sensor of Linear Moving Magnet Compressor's PistonStroke, IEEE Sensors Journal, vol. 9, n. 2, p. 154-158, February 2009).This last method only requires knowledge of electrical parameters, andno information about the mechanical subsystem of the actuator areneeded. However, coil current derivative must be computed which is nottrivial in a noisy environment.

For example, one method presented a velocity observer to estimate theback EMF that does not rely on computing any time. (See, e.g., J.Latham, et al., Parameter Estimation and a Series of Nonlinear Observersfor the System Dynamics of a Linear Vapor Compressor, IEEE Transactionson Industrial Electronics, vol. 63, no 11, p. 6736-6744, November 2016).The resulting position from integrating the estimated velocity issensitive to measurement bias that propagates into the velocityestimation which results in drift when integrated. This effect can bebounded by adding another stage to the observer. (See, e.g., P.Mercorelli, A Motion-Sensorless Control for Intake Valves in CombustionEngines, IEEE Transactions on Industrial Electronics, vol. 64, n 4, p.3402-3412, April 2017). This additional stage adds partial knowledgeabout the mechanical subsystem of the actuator, and is robust tounknown, bounded forces. However, these studies are limited to a lineardomain of the actuator, where the parameters of the equivalent electriccircuit of the actuator can be approximated as constants, which is notvalid for vibrating membrane pumps where the actuator is made as smallas possible.

In view of the foregoing, there exists a need for controlling an energyefficient implantable pump that has light weight, small size, and faststart and stop response, for example, without relying on position,velocity, or acceleration sensors.

It would further be desireable to provide an improved controller forcontrolling an energy efficient implantable pump relying on positionmeasurement.

SUMMARY OF THE INVENTION

The present invention overcomes the drawbacks of previously-known LVADsystems and methods by providing an implantable pump system having anundulating membrane capable of producing a wide range of physiologicalflow rates while applying low shear forces to the blood, therebyreducing hemolysis and platelet activation relative to previously-knownsystems.

In accordance with one aspect of the invention, the implantable bloodpump system includes an implantable blood pump configured to beimplanted at a patient's heart, and a controller operatively coupled tothe implantable blood pump. The implantable blood pump includes ahousing having an inlet and an outlet, a deformable membrane disposedwithin the housing, and an actuator having a stationary component and amoving component coupled to the deformable membrane. The actuator ispowered by an alternating current that causes the moving component toreciprocate at a predetermined frequency and amplitude relative to thestationary component, thereby causing the deformable membrane to producea predetermined blood flow from the inlet out through the outlet.

In addition, the controller is programmed to operate the actuator tocause the moving component to reciprocate at the predetermined frequencyand amplitude relative to the stationary component, receive a signalindicative of the alternating current via a current sensor operativelycoupled to the controller, determine a position of the moving componentbased on the signal indicative of the alternating current, and adjustoperation of the actuator to cause the moving component to reciprocateat an adjusted predetermined frequency and amplitude relative to thestationary component based on the position of the moving component,thereby causing the deformable membrane to produce an adjustedpredetermined blood flow from the inlet out through the outlet. Forexample, the adjusted predetermined blood flow may be a pulsesynchronized with the patient's heartbeat.

The controller may be programmed to determine the position of the movingcomponent by estimating a velocity of the moving component based on thesignal indicative of the alternating current. For example, thecontroller may be programmed to estimate the velocity of the movingcomponent based on co-energy W values of a finite elements model (FEM)of various positions and alternating currents of the moving component.In addition, the controller may be programmed to determine the positionof the moving component by determining the velocity of the movingcomponent based on the estimated velocity of the moving component.

Further, the controller may be programmed to adjust operation of theactuator to cause the moving component to reciprocate at the adjustedpredetermined frequency and amplitude relative to the stationarycomponent while limiting overshoot. For example, the controller mayinclude a proportional integral (PI) controller programmed to limitovershoot by canceling errors due to un-modeled dynamics of theimplantable blood pump. The controller may be programmed to determinethe position of the moving component based on the signal indicative ofthe alternating current and variations of inductance and back EMFcoefficient.

In accordance with one aspect of the present invention, the stationarycomponent includes an electromagnetic assembly that generates a magneticfield. Moreover, the moving component may include a magnetic ringconcentrically suspended around the electromagnetic assembly anddesigned to reciprocate responsive to the magnetic field at thepredetermined frequency and amplitude over the electromagnetic assembly.The electromagnetic assembly may include first and secondelectromagnetic coils, such that the magnetic ring is caused to movewhen at least one of the first or second electromagnetic coils ispowered by the alternating current. In addition, the magnetic ringinduces wave-like deformations in the deformable membrane byreciprocating over the electromagnetic assembly.

In addition, the implantable blood pump may include first and secondsuspension rings concentrically disposed around and coupled to thestationary component and the moving component. Accordingly, the movingcomponent may be coupled to each of the deformable membrane and thefirst and second suspension rings via a plurality of posts, such thatthe first and second suspension rings permit the moving component toreciprocate relative to the stationary component. The first and secondsuspension rings may exert a spring force on the moving component whenthe moving component reciprocates relative to the stationary component.

Additionally, the implantable blood pump further may include a rigidring coupled to the moving component and to the deformable membrane.Moreover, a bottom surface of the actuator and an interior portion ofthe housing adjacent the outlet may form a flow channel within which thedeformable membrane is suspended. Accordingly, the deformable membranemay have a central aperture adjacent the outlet. In addition, theactuator and an interior surface of the housing adjacent the inlet mayform a delivery channel extending from the inlet to the flow channel.The implantable blood pump system further may include a rechargeablebattery for delivering the alternating current to power the implantableblood pump.

In accordance with another aspect of the present invention, analternative exemplary implantable blood pump system is provided. Thesystem may include the implantable blood pump sized and shaped to beimplanted at a patient's heart described above, and a controlleroperatively coupled to the implantable blood pump. For example, thecontroller may be programmed to: operate the actuator to cause themoving component to reciprocate at the predetermined frequency andamplitude relative to the stationary component; receive a signalindicative of an intensity of a magnetic field of a magnet coupled tothe moving component via a sensor, e.g., a hall effector sensor,operatively coupled to the controller, the sensor stationary relative tothe stationary component; determine a position of the moving componentbased on the signal indicative of the intensity of the magnetic field;and adjust operation of the actuator to cause the moving component toreciprocate at an adjusted predetermined frequency and amplituderelative to the stationary component based on the position of the movingcomponent, thereby causing the deformable membrane to produce anadjusted predetermined blood flow from the inlet out through the outlet.

For example, the sensor may be coupled to the stationary component orthe housing. The controller further may be programmed to estimate bloodflow from the inlet out through the outlet based on the position of themoving component. Additionally, the controller further may be programmedto detect a fault by comparing an average residual value based on theposition of the moving component with a predetermined threshold value.

In accordance with another aspect of the present invention, thecontroller may be programmed to: operate the actuator to cause themoving component to reciprocate at the predetermined frequency andamplitude relative to the stationary component; receive a signalindicative of an intensity of a magnetic field of a magnet coupled tothe moving component via a sensor operatively coupled to the controller,the sensor stationary relative to the stationary component; estimate avelocity of the moving component based on the signal indicative of theintensity of the magnetic field; and adjust operation of the actuator tocause the moving component to reciprocate at an adjusted predeterminedfrequency and amplitude relative to the stationary component based onthe velocity of the moving component, thereby causing the deformablemembrane to produce an adjusted predetermined blood flow from the inletout through the outlet.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts an exemplary embodiment of the pump system of the presentinvention comprising an implantable pump, controller, battery,programmer and mobile device.

FIG. 2 is a perspective view of the implantable pump of FIG. 1.

FIGS. 3A and 3B are, respectively, a perspective view and a schematicview of the electronic components of an exemplary embodiment of thecontroller of the present invention.

FIG. 4 is a plan view of an extracorporeal battery for use in the pumpsystem of the present invention.

FIGS. 5A and 5B are, respectively, a perspective view and a schematicview of the electronic components of an exemplary embodiment of theprogrammer of the present invention.

FIG. 6 is a perspective view of the pump assembly of the presentinvention.

FIG. 7 is a perspective, cut-away view of the implantable pump of thepresent invention.

FIG. 8 is an exploded view of the implantable pump of the presentinvention.

FIG. 9 is a perspective cross-sectional view of the pump assembly of thepresent invention.

FIG. 10 is a perspective cross-sectional view of the membrane assemblyof the present invention.

FIG. 11 is a perspective cross-sectional view of the moving componentsof the pump assembly according to a first embodiment of the presentinvention.

FIG. 12 is a cross-sectional view of the implantable pump of the presentinvention.

FIG. 13 is a cross-sectional view of a lower portion of the implantablepump depicting the flow channel and membrane assembly in a restingposition.

FIG. 14 is a cross-sectional view of a lower portion of the implantablepump depicting the flow channel and membrane assembly with the membraneundulating.

FIG. 15A is a cross-sectional view of an alternative exemplaryembodiment of an implantable pump of the present invention with improvedhydraulic performance for use in the pump system of FIG. 1.

FIG. 15B is a perspective view of the implantable pump of FIG. 15A.

FIG. 16A illustrates blood flow across a planar ring membrane support,whereas FIG. 16B illustrates blood flow using a pump assembly with askirt in accordance with one aspect of the present invention.

FIG. 17 shows graphs illustrating the relationship between max hydraulicpower and the height of the skirt.

FIG. 18 is a cross-sectional view of yet another alternative exemplaryembodiment of an implantable pump of the present invention with improvedhydraulic performance, wherein the outflow cannula is disposed coaxiallywithin the inflow cannula.

FIG. 19 is a cross-sectional view of yet another alternative exemplaryembodiment of an implantable pump of the present invention having a ringand skirt with improved hydraulic performance for use in the pump systemof FIG. 1.

FIG. 20 is a cross-sectional view of yet another alternative exemplaryembodiment of an implantable pump of the present invention having aring, skirt, and expandable portion with improved hydraulic performancefor use in the pump system of FIG. 1.

FIGS. 21A-H illustrates various configurations for coupling a battery toa controller of the present invention, and FIG. 21I illustrates acontroller coupled to a power supply.

FIG. 22 is a flow chart illustrating steps of an exemplary method forcontrolling an implantable pump constructed in accordance with theprinciples of the present invention.

FIG. 23 is a FEM model of a subset of an actuator assembly of animplantable pump constructed in accordance with the principles of thepresent invention.

FIG. 24A displays co-energy as a function of magnetic ring position andcoil current, FIG. 24B displays force as a function of magnetic ringposition and coil current, FIG. 24C displays circuit inductance of anequivalent circuit as a function of magnetic ring position and coilcurrent, and FIG. 24D displays circuit EMF coefficient of an equivalentcircuit as a function of magnetic ring position and coil current.

FIG. 25 is a graph illustrating spring reaction force of as a functionof position of an implantable pump.

FIG. 26 is a schematic of the parameters of an equivalent circuit inaccordance with the principles of the present invention.

FIGS. 27A-D is a schematic of the power electronics constructed inaccordance with the principles of the present invention.

FIG. 28 illustrates ADC sampling for measuring current in accordancewith the principles of the present invention.

FIG. 29A is a diagram illustrating multistage control of a controllerconstructed in accordance with the principles of the present invention,and FIG. 29B is a diagram illustrating an alternative multistage controlof a controller constructed in accordance with the principles of thepresent invention.

FIGS. 30A-C illustrate identification of variations of resistance,inductance, and EMF coefficient, respectively, with magnetic ringposition and coil current in accordance with the principles of thepresent invention.

FIGS. 31A and 31B illustrate the system response to a desired strokewith respect to coil current and magnetic ring position, respectively.

FIGS. 32A and 32B illustrate the system response to change of operationpoints with respect to amplitude and frequency, respectively.

FIGS. 33A-D illustrate stroke output error maps of the system inaccordance with the principles of the present invention.

FIG. 34 is a diagram illustrating multistage control of a controllerrelying on position measurement constructed in accordance with theprinciples of the present invention.

FIG. 35 is a diagram illustrating an alternative multistage sensorlesscontrol of a controller constructed in accordance with the principles ofthe present invention.

FIG. 36 is a diagram illustrating yet another alternative multistagesensorless control of a controller constructed in accordance with theprinciples of the present invention.

DETAILED DESCRIPTION

The implantable pump system of the present invention is particularlywell-suited for use as a left ventricular assist device (LVAD), andincludes an undulating membrane pump suitable for long-term implantationin a patient having end term heart failure. An implantable pump systemconstructed in accordance with the principles of the present inventionincludes an implantable pump and an extracorporeal battery, controllerand programmer. The implantable pump includes a housing having an inlet,and outlet, a flexible membrane, and an actuator assembly. Whenconfigured as an LVAD, the housing includes an inlet cannula that isinserted into a patient's left ventricle near the apex and an outletcannula that is surgically placed in fluid communication with thepatient's aorta. By activating the actuator assembly within theimplantable pump, membrane is induced to undulate, thereby causing bloodto be drawn into the pump through the inlet cannula and expelled throughthe outlet cannula into the aorta. Flow rate and pulsatility may bemanipulated by changing one or more of the frequency, amplitude and dutycycle of the actuator assembly.

For improved hydraulic performance, the implantable pump may include askirt disposed within the housing to guide blood flow from the inlet ofthe pump towards the outlet. The skirt may be positioned within thehousing such that blood flows across opposing sides of the skirt andtowards the undulating membrane upon activation of the pump.

Referring now to FIG. 1, pump system 10 constructed in accordance withthe principles of the present invention is described. Pump system 10includes implantable pump 20, controller 30, battery 40, programmer 50and optionally, a software module programmed to run on mobile device 60.Implantable pump 20 is configured to be implanted within a patient'schest so that inlet cannula 21 is coupled to left ventricle LV of heartH. Outlet cannula 22 of pump 20 is configured to be coupled to aorta A.Inlet cannula 21 preferably is coupled to the apex of left ventricle LV,while outlet cannula 22 is coupled to aorta A in the vicinity of theascending aorta, above the level of the cardiac arteries. Implantablepump 20 may be affixed within the patient's chest using a ring-suture orother conventional technique. Outlet cannula 22, which may comprise aDacron graft or other synthetic material, is coupled to outlet 23 ofimplantable pump 20.

Referring now also to FIG. 2, implantable pump 20 in a preferredembodiment consists of upper housing portion 24 joined to lower housingportion 25 along interface 26, for example, by threads or welding, toform fluid tight pump housing 27 that may have a cylindrical shape.Upper housing portion 24 includes inlet cannula 21 and electricalconduit 28 for receiving electrical wires from controller 30 and battery40. Lower housing portion 25 includes outlet 23 that couples to outletcannula 22, as shown in FIG. 1. Pump housing 27 is made of abiocompatible material, such as stainless steel, and is sized to beimplanted within a patient's chest.

Referring again to FIG. 1, in one embodiment, controller 30 and battery40 are extracorporeal, and are sized so as to be placed on a belt orgarment worn by the patient. Both controller 30 and battery 40 areelectrically coupled to implantable pump 20, for example, via cable 29that extends through a transcutaneous opening in the patient's skin andinto electrical conduit 28 of pump housing 27. Illustratively, battery40 is electrically coupled to controller 30 via cable 41 that isintegrated into belt 42. In an alternative embodiment, controller 30 maybe enclosed within a biocompatible housing and sized to be implantedsubcutaneously in the patient's abdomen. In this alternative embodiment,controller 30 may include a wireless transceiver for bi-directionalcommunications with an extracorporeal programming device and alsoinclude a battery that is continuously and inductively charged viaextracorporeal battery 40 and an extracorporeal charging circuit. Aswill be understood, the foregoing alternative embodiment avoids the useof transcutaneous cable 29, and thus eliminates a frequent source ofinfection for conventional LVAD devices.

Battery 40 preferably comprises a rechargeable battery capable ofpowering implantable pump 20 and controller 30 for a period of severaldays, e.g., 3-5 days, before needing to be recharged. Battery 40 mayinclude a separate charging circuit, not shown, as is conventional forrechargeable batteries. Battery 40 preferably is disposed within ahousing suitable for carrying on a belt or holster, so as not tointerfere with the patient's daily activities.

Programmer 50 may consist of a conventional laptop computer that isprogrammed to execute programmed software routines, for use by aclinician or medical professional, for configuring and providingoperational parameters to controller 30. The configuration andoperational parameter data is stored in a memory associated withcontroller 30 and used by the controller to control operation ofimplantable pump 20. As described in further detail below, controller 30directs implantable pump 20 to operate at specific parameters determinedby programmer 50. Programmer 50 preferably is coupled to controller 30via cable 51 only when the operational parameters of the implantablepump are initially set or periodically adjusted, e.g., when the patientvisits the clinician.

In accordance with another aspect of the invention, mobile device 60,which may a conventional smartphone, may include an application programfor bi-directionally and wirelessly communicating with controller 30,e.g., via WiFi or Bluetooth communications. The application program onmobile device 60 may be programmed to permit the patient to sendinstructions to controller to modify or adjust a limited number ofoperational parameters of implantable pump 20 stored in controller 30.Alternatively or in addition, mobile device 60 may be programmed toreceive from controller 30 and to display on screen 61 of mobile device60, data relating to operation of implantable pump 20 or alert or statusmessages generated by controller 30.

With respect to FIGS. 3A and 3B, controller 30 is described in greaterdetail. As depicted in FIG. 1, controller 30 may be sized and configuredto be worn on the exterior of the patient's body and may be incorporatedinto a garment such as a belt or a vest. Controller 30 includes inputport 31, battery port 32, output port 33, indicator lights 34, display35, status lights 36 and buttons 37.

Input port 31 is configured to periodically and removably accept cable51 to establish an electrical connection between programmer 50 andcontroller 30, e.g., via a USB connection. In this manner, a clinicianmay couple to controller 30 to set or adjust operational parametersstored in controller 30 for controlling operation of implantable pump.In addition, when programmer 50 is coupled to controller 30, theclinician also may download from controller 30 data relating tooperation of the implantable pump, such as actuation statistics, forprocessing and presentation on display 55 of programmer 50.Alternatively, or in addition, controller 30 may include a wirelesstransceiver for wirelessly communicating such information withprogrammer 50. In this alternative embodiment, wireless communicationsbetween controller 30 and programmer 50 may be encrypted with anencryption key associated with a unique identification number of thecontroller, such as a serial number.

Battery port 32 is configured to removably accept cable 41,illustratively shown in FIG. 1 as integrated with belt 42, so that cable41 routed through the belt and extends around the patient's back untilit couples to controller 30. In this manner, battery 40 may be removedfrom belt 42 and disconnected from controller 30 to enable the patientto periodically replace the battery with a fully charged battery. It isexpected that the patient will have available to him or her at least twobatteries, so that while one battery is coupled to controller 30 toenergize the controller and implantable pump, the other battery may beconnected to a recharging station. Alternatively, or in addition,battery port 32 may be configured to accept a cable that is coupleddirectly to a power supply, such a substantially larger battery/chargercombination that permits the patient to remove battery 40 while lyingsupine in a bed, e.g., to sleep.

Output port 33 is electrically coupled to cable 29, which in turn iscoupled to implantable pump 20 through electrical conduit 28 of pumphousing 27. Cable 29 provides both energy to energize implantable pump20 in accordance with the configuration settings and operationalparameters stored in controller 30, and to receive data from sensorsdisposed in implantable pump 20. In one embodiment, cable 29 maycomprise an electrical cable having a biocompatible coating and isdesigned to extend transcutaneously. Cable 29 may be impregnated withpharmaceuticals to reduce the risk of infection, the transmission ofpotentially hazardous substances or to promote healing where it extendsthrough the patient's skin.

As mentioned above, controller 30 may include indicator lights 34,display 35, status lights 36 and buttons 37. Indicator lights 34 mayvisually display information relevant to operation of the system, suchas the remaining life of battery 40. Display 35 may be a digital liquidcrystal display that displays real time pump performance data,physiological data of the patient, such as heart rate, or operationalparameters of the implantable pump, such as the target pump pressure orflow rate, etc. When it is determined that certain parameter conditionsexceed preprogrammed thresholds, an alarm may be sounded and an alertmay be displayed on display 35. Status lights 36 may comprise lightemitting diodes (LEDs) that are turned on or off to indicate whethercertain functionality of the controller or implantable pump is active.Buttons 37 may be used to wake up display 35, to set or quiet alarms,etc.

With respect to FIG. 3B, the components of the illustrative embodimentof controller 30 of FIG. 3A are described. In addition to the componentsof controller 30 described in connection with FIG. 3A, controller 30further includes microprocessor 38, memory 39, battery 43, optionaltransceiver 44 and amplifier circuitry 45. Microprocessor may be ageneral purpose microprocessor, for which programming to controloperation of implantable pump 20 is stored in memory 39. Memory 39 alsomay store configuration settings and operational parameters forimplantable pump 20. Battery 40 supplies power to controller 30 toprovide continuity of operation when battery 40 is periodically swappedout. Optional transceiver 44 to facilitates wireless communication withprogrammer 50 and/or mobile device 60 via any of a number of well-knowncommunications standards, including BLUETOOTH™, ZigBee, and/or any IEEE802.11 wireless standard such as Wi-Fi or Wi-Fi Direct. Controller 30further may include amplifier circuitry 45 for amplifying electricalsignals transferred between controller 30 and implantable pump 20.

Referring now to FIG. 4, battery 40 is described. Battery 40 providespower to implantable pump 20 and also may provide power to controller30. Battery 40 may consist of a single battery or a plurality ofbatteries disposed within a housing, and preferably is sized andconfigured to be worn on the exterior of the patient's body, such as onbelt 42. Battery life indicator 46 may be provided on the exterior ofbattery 40 to indicate the degree to the remaining charge of thebattery. Cable 41 may have one end removably coupled to battery 40 andthe other end removably coupled to battery port of controller 30 tosupply power to energize implantable pump 20. In one embodiment, battery40 may be rechargeable using a separate charging station, as is known inthe art of rechargeable batteries. Alternatively, or in addition,battery 40 may include port 47 which may be removably coupled to atransformer and cable to permit the battery to be recharged using aconventional residential power outlet, e.g., 120 V, 60 Hz AC power.

Referring now to FIGS. 5A-5B, programmer 50 is described. Programmer 50may be conventional laptop loaded with programmed software routines forconfiguring controller 30 and setting operational parameters thatcontroller 30 uses to control operation of implantable pump 20. Asdiscussed above, programmer 50 typically is located in a clinician'soffice or hospital, and is coupled to controller 30 via cable 51 orwirelessly to initially set up controller 30, and then periodicallythereafter as required to adjust the operational parameters as may beneeded. The operation parameters of controller 30 set using theprogrammed routines of programmer 50 may include but are not limited toapplied voltage, pump frequency, pump amplitude, target flow rate,pulsatility, etc. When first implanted, the surgeon or clinician may useprogrammer 50 to communicate initial operating parameters to controller30. Following implantation, the patient periodically may return to theclinician's office for adjustments to the operational parameters whichmay again be made using programmer 50.

Programmer 50 may be any type of conventional personal computer devicesuch as a laptop or a tablet computer having touch screen capability. Asillustrated in FIG. 5B, programmer 50 preferably includes processor 52,memory 53, input/output device 54, display 55, battery 56 andcommunication unit 57. Memory 53 may include the operating system forthe programmer, as well as the programmed routines needed to communicatewith controller 30. Communication unit 57 may include any of a number ofwell-known communication protocols, such as BLUETOOTH™, ZigBee, and/orany IEEE 802.11 wireless standard such as Wi-Fi or Wi-Fi Direct. Asillustrated in FIG. 5A, the programmed routines used to program andcommunicate with controller 30 also may provide data for display on thescreen of programmer 50 identifying operational parameters with whichcontroller 30 controls implantable pump 20. The programmed routines alsomay enable programmer 50 to download from controller 30 operational dataor physiologic data communicated by the implantable pump and to displaythat information in real time while the programmer is coupled to thecontroller via a wired or wireless connection. The transferred data maythen be processed and displayed on the screen of programmer 50.

Referring now to FIGS. 6 and 7, a preferred embodiment of pump assembly70 and implantable pump 20 are illustrated. However, it is understoodthat pump assemblies and implantable pumps, and components includedtherein, may have different shapes and sizes than those illustrated inFIGS. 6 and 7 without departing from the invention described herein. Asis illustrated in FIG. 7, pump assembly 70 is configured to fit withinpump housing 27. To fix pump assembly 70 within pump housing 27, pumpassembly 70 may include fixation ring 71, which may extend from andaround stator assembly 72, and may be captured between upper housingportion 24 and lower housing portion 25 when the housing portions areassembled, as illustrated in FIG. 7. In this manner, stator assembly 72may be suspended within the pump housing in close-fitting relation tothe interior walls of the pump housing. Fixation ring 71 preferably is arigid annular structure that is disposed concentrically around statorassembly 72, having a larger diameter than stator assembly 72. Fixationring 71 may be rigidly coupled to stator assembly 72 via struts 73.Struts 73 may create gap 74 between fixation ring 71 and stator assembly72, which preferably is about 0.05 mm at its most restricted point.

As shown in FIG. 7, pump assembly 70 may be disposed in pump housing 27such that fixation ring 71 is captured on step 75 formed between upperhousing portion 24 and lower housing portion 25. In this manner, statorassembly 72 may be suspended within, and prevented from moving within,pump housing 27. Pump housing 27 preferably is sized and configured toconform to pump assembly 70 such that, stator assembly 72 does notcontact the interior of the pump housing at any location other than atfixation ring 71.

FIG. 8 is an exploded view of implantable pump 20, depicting thearrangement of the internal components of pump assembly 70 arrangedbetween upper housing portion 24 and lower housing portion 25. Inparticular, pump assembly 70 may comprise stator assembly 72, magneticring assembly 76, first electromagnetic coil 77, second electromagneticcoil 78, fixation ring 71, first suspension ring 79, second suspensionring 80, posts 81 and membrane assembly 82. Stator assembly 72 maycomprise tapered section 83, electromagnetic coil holder portions 84, 85and 86, and flanged portion 87. Magnetic ring assembly 76 may comprisemagnetic ring 88 and magnetic ring holder portions 89 and 90. First andsecond electromagnetic coils 77 and 78, together with electromagneticcoil holder portions 84, 85 and 86 may form electromagnet assembly 91.Electromagnet assembly 91 together with stator assembly 72 form anactuator assembly. The actuator assembly together with magnetic ringassembly 76 in turn forms the actuator system of implantable pump 20.

First electromagnetic coil 77 and second electromagnetic coil 78 may beconcentrically sandwiched between electromagnetic coil holder portions84, 85 and 86 to form electromagnet assembly 91. Tapered section 83,which may be coupled to fixation ring 71 and first suspension spring 79,may be located concentrically atop electromagnet assembly 91. Magneticring 88 may be disposed with magnetic ring holder portions 89 and 90 toform magnetic ring assembly 76, which may be concentrically disposed forreciprocation over electromagnet assembly 91. Second suspension ring 80may be disposed concentrically beneath electromagnet assembly 91.Flanged portion 87 may be concentrically disposed below secondsuspension ring 80. Posts 81 may engage first suspension ring 79,magnetic ring assembly 76 and second suspension ring 80 at equallyspaced locations around the actuator assembly. Membrane assembly 82 maybe positioned concentrically below flanged portion 87 and engaged withposts 81.

Further details of pump assembly 70 are provided with respect to FIG. 9.Specifically, actuator assembly 95 comprises stator assembly 72 andelectromagnet assembly 91, including first and second electromagneticcoils 77 and 78. During use of implantable pump 20, actuator assembly 95remains stationary relative to pump housing 27. First electromagneticcoil 77 and second electromagnetic coil 78 may be separated byelectromagnetic holder portion 85. Controller 30 and battery 40 areelectrically coupled to electromagnetic coils 77 and 78 via cable 29that extends through electrical conduit 28 of pump housing 27 to supplycurrent to electromagnetic coils 77 and 78. First electromagnetic coil77 and second electromagnetic coil 78 may be in electrical communicationwith one another or may be configured to operate independently and haveseparate wired connections to controller 30 and battery 40 via cable 29.

Electromagnetic coils 77 and 78 may be made of any electricallyconductive metallic material such as copper and further may comprise ofone or more smaller metallic wires wound into a coil. The wires of theelectromagnetic coils are insulated to prevent shorting to adjacentconductive material. Other components of pump assembly 70, such asstator assembly 72, preferably also are insulated and/or made ofnon-conductive material to reduce unwanted transmission of theelectrical signal.

Actuator assembly 95 may be surrounded by first suspension ring 79 andsecond suspension ring 80. Suspension rings 79 and 80 may be annular inshape and fit concentrically around actuator assembly 95. Firstsuspension ring 79 preferably is rigidly affixed to tapered section 83near a top portion of stator assembly 72 via struts 73 extending fromthe suspension ring to the stator assembly. As discussed above, struts73 may also affix fixation ring 71 to stator assembly 72. Fixation ring71 and first suspension spring 79 may be sized and positioned such thata gap of no less than 0.5 mm exists between first suspension ring 79 andfixation ring 71. Second suspension ring 80 similarly may be rigidlyaffixed via struts near the bottom of stator assembly 72, belowelectromagnet assembly 91. Suspension rings 79 and 80 preferably aresized and shaped such that when suspension rings 79 and 80 arepositioned surrounding actuator assembly 95, a gap of no less than 0.5mm exists between actuator assembly 95 and suspension rings 79 and 80.

First suspension ring 79 and second suspension ring 80 may comprisestainless steel having elastic properties and which exhibits a springforce when deflected in a direction normal to the plane of the spring.First suspension ring 79 and second suspension ring 80 may besubstantially rigid with respect to forces applied tangential to thesuspension ring. In this manner, first suspension ring 79 and secondsuspension ring 80 may exhibit a spring tension when deformed up anddown relative to a vertical axis of the actuator assembly but mayrigidly resist movement along any other axis, e.g., tilt or twistmovements.

Magnetic ring assembly 76 may be annular in shape and concentricallysurrounds actuator assembly 95. Magnetic ring 88 may comprise one ormore materials exhibiting magnetic properties such as iron, nickel,cobalt or various alloys. Magnetic ring 88 may be made of a singleunitary component or comprise several magnetic components that arecoupled together. Magnetic ring assembly 76 may be sized and shaped suchthat when it is positioned concentrically over actuator assembly 95, agap of no less than 0.5 mm exists between an outer lateral surface ofactuator assembly 95 and an interior surface of magnetic ring assembly76.

Magnetic ring assembly 76 may be concentrically positioned aroundactuator assembly 95 between first suspension ring 79 and secondsuspension ring 80, and may be rigidly coupled to first suspension ring79 and second suspension ring 80. Magnetic ring assembly 76 may berigidly coupled to the suspension rings by more than one post 81 spacedevenly around actuator assembly 95 and configured to extend parallel toa central axis of pump assembly 70. Suspension rings 79 and 80 andmagnetic ring assembly 76 may be engaged such that magnetic ringassembly 76 is suspended equidistant between first electromagnetic coil77 and second electromagnetic coil 78 when the suspension rings are intheir non-deflected shapes. Each of suspension rings 79 and 80 andmagnetic ring holder portions 89 and 90 may include post receivingregions for engaging with posts 81 or may be affixed to posts 81 in anysuitable manner that causes suspension rings 79 and 80 and magnetic ringassembly 76 to be rigidly affixed to posts 81. Posts 81 may extendbeyond suspension rings 79 and 80 to engage other components, such asflanged portion 87 and membrane assembly 82.

First electromagnetic coil 77 may be activated by controller applying anelectrical signal from battery 40 to first electromagnetic coil 77, thusinducing current in the electromagnetic coil and generating a magneticfield surrounding electromagnetic coil 77. The direction of the currentin electromagnetic coil 77 and the polarity of magnetic ring assembly 76nearest electromagnetic coil 77 may be configured such that the firstelectromagnetic coil magnetically attracts or repeals magnetic ringassembly 76 as desired. Similarly, a magnetic field may be created insecond electromagnetic coil 78 by introducing a current in the secondelectromagnetic coil. The direction of the current in secondelectromagnetic coil 78 and the polarity of magnetic ring assembly 76nearest the second electromagnetic coil also may be similarly configuredso that first electromagnetic coil 77 magnetically attracts or repelsmagnetic ring assembly 76 when an appropriate current is induced insecond electromagnetic coil 78.

Because magnetic ring assembly 76 may be rigidly affixed to posts 81,which in turn may be rigidly affixed to first suspension ring 79 andsecond suspension ring 80, the elastic properties of the suspensionrings permit magnetic ring assembly 76 to move up towards firstelectromagnetic coil 77 or downward toward second electromagnetic coil78, depending upon the polarity of magnetic fields generated by theelectromagnetic rings. In this manner, when magnetic ring assembly 76experiences an upward magnetic force, magnetic ring assembly 76 deflectsupward towards first electromagnetic coil 77. As posts 81 move upwardwith magnetic ring assembly 76, posts 81 cause the suspensions rings 79and 80 to elastically deform, which creates a spring force opposite tothe direction of movement. When the magnetic field generated by thefirst electromagnetic coil collapses, when the electrical currentceases, this downward spring force causes the magnetic ring assembly toreturn to its neutral position. Similarly, when magnetic ring assembly76 is magnetically attracted downward, magnetic ring assembly 76deflects downward towards second electromagnetic ring 78. As posts 81move downward with magnetic ring assembly 76, posts 81 impose an elasticdeformation of the first and second suspension rings, thus generating aspring force in the opposite direction. When the magnetic fieldgenerated by the second electromagnetic ring collapses, when theelectrical current ceases, this upward spring force causes the magneticring assembly to again return to its neutral position.

Electromagnetic coils 77 and 78 may be energized separately, oralternatively, may be connected in series to cause the electromagneticcoils to be activated simultaneously. In this configuration, firstmagnetic coil may be configured to experience a current flow directionopposite that of the second electromagnetic coil. Accordingly, whencurrent is induced to first electromagnetic coil 77 to attract magneticring assembly 76, the same current is applied to second electromagneticcoil 78 to induce a current that causes second electromagnetic coil 78to repel magnetic ring assembly 76. Similarly, when current is inducedto second electromagnetic coil 78 to attract magnetic ring assembly 76,the current applied to first electromagnetic coil 77 causes the firstelectromagnetic coil to repel magnetic ring assembly 76. In this manner,electromagnetic coils 77 and 78 work together to cause deflection ofmagnetic ring assembly 76.

By manipulating the timing and intensity of the electrical signalsapplied to the electromagnetic coils, the frequency at which magneticring assembly 76 deflects towards the first and second electromagneticcoils may be altered. For example, by alternating the current induced inthe electromagnetic coils more frequently, the magnetic ring assemblymay be caused to cycle up and down more times in a given period. Byincreasing the amount of current, the magnetic ring assembly may bedeflected at a faster rate and caused to travel longer distances.

Alternatively, first electromagnetic coil 77 and second electromagneticcoil 78 may be energized independently. For example, firstelectromagnetic coil 77 and second electromagnetic coil 78 may beenergized at varying intensities; one may be coordinated to decreaseintensity as the other increases intensity. In this manner, intensity ofthe signal applied to second electromagnetic coil 78 to cause downwardmagnetic attraction may simultaneously be increased as the intensity ofthe signal applied to first electromagnetic coil 77 causes an upwardmagnetic attraction that decreases.

In accordance with one aspect of the invention, movements of magneticring assembly 76 may be translated to membrane assembly 82 which may bedisposed concentrically below stator assembly 72. Membrane assembly 82preferably is rigidly attached to magnetic ring assembly 76 by posts 81.In the embodiment depicted in FIG. 9, posts 81 may extend beyond secondsuspension ring 80 and coupled to membrane assembly 82.

Referring now to FIG. 10, one embodiment of membrane assembly 82 isdescribed in greater detail. Membrane assembly 82 may comprise rigidmembrane ring 96 and membrane 97. Rigid membrane ring 96 exhibits rigidproperties under typical forces experienced during the full range ofoperation of the present invention. Post reception sites 98 may beformed into rigid membrane ring 96 to engage membrane assembly 82 withposts 81. Alternatively, posts 81 may be attached to rigid membrane ring96 in any other way which directly translates the motion of magneticring assembly 76 to rigid membrane ring 96. Rigid membrane ring 96 maybe affixed to membrane 97 and hold the membrane in tension. Membrane 97may be molded directly onto rigid membrane ring 96 or may be affixed torigid membrane ring 96 in any way that holds membrane 97 uniformly intension along its circumference. Membrane 97 alternatively may include aflexible pleated structure where it attaches to rigid membrane ring 96to increase the ability of the membrane to move where the membrane isaffixed to rigid membrane ring 96. Membrane 97 may further includecircular aperture 99 disposed in the center of the membrane.

In a preferred embodiment, membrane 97 has a thin, planar shape and ismade of an elastomer having elastic properties and good durability.Alternatively, membrane 97 may have a uniform thickness from themembrane ring to the circular aperture. As a yet further alternative,membrane 97 may vary in thickness and exhibit more complex geometries.For example, as shown in FIG. 10, membrane 97 may have a reducedthickness as the membrane extends from rigid membrane ring 96 tocircular aperture 99. Alternatively, or in addition to, membrane 97 mayincorporate metallic elements such as a spiral spring to enhance thespring force of the membrane in a direction normal to plane of themembrane, and this spring force may vary radially along the membrane. Inyet another embodiment, membrane 97 may be pre-formed with an undulatingshape.

FIG. 11 depicts moving portions of the embodiment of pump assembly 70shown in FIGS. 6-9 as non-grayed out elements. Non-moving portions ofthe pump assembly, including actuator assembly 95 and electromagnetassembly 91 (partially shown) may be fixed to pump housing 27 byfixation ring 71. Moving portions of pump assembly 70 may include posts81, first suspension spring 79, magnetic ring assembly 76, secondsuspension spring 80 and membrane assembly 82. As magnetic ring assembly76 moves up and down, the movement is rigidly translated by posts 81 tomembrane assembly 82. Given the rigidity of the posts, when magneticring assembly 76 travels a certain distance upward or downward, membraneassembly 82 may travel the same distance. For example, when magneticring assembly 76 travels 2 mm from a position near first electromagneticcoil 77 to a position near second electromagnetic coil 78, membraneassembly 82 may also travel 2 mm in the same direction. Similarly, thefrequency at which magnetic ring assembly 76 traverses the space betweenthe first and second electromagnetic coils may be the same frequency atwhich membrane assembly 82 travels the same distance.

Referring now to FIG. 12, in the embodiment of implantable pump 20described in FIGS. 6-9, blood may enter implantable pump 20 from theleft ventricle through inlet cannula 21 and flow downward along pumpassembly 70 into delivery channel 100, defined by the interior surfaceof pump housing 27 and exterior of pump assembly 70. Delivery channel100 begins at the top of stator assembly 72 and extends between taperedsection 83 and the interior of pump housing 27. As the blood moves downtapered section 83, it is directed through gap 74 and into a verticalportion of delivery channel 100 in the area between pump housing 27 andactuator assembly 95, and including in the gap between magnetic ringassembly 76 and electromagnet assembly 91. Delivery channel 100 extendsdown to flanged portion 87 of stator assembly 72, which routes bloodinto flow channel 101, within which membrane assembly 82 is suspended.By directing blood from inlet cannula 21 through delivery channel 100 toflow channel 101, delivery channel 100 delivers blood to membraneassembly 82. By actuating electromagnetic coils 77 and 78, membrane 97may be undulated within flow channel 101 to induce wavelike formationsin membrane 97 that move from the edge of the membrane towards circularaperture 99. Accordingly, when blood is delivered to membrane assembly82 from delivery channel 100, it may be propelled radially along boththe top and bottom of membrane 97 towards circular aperture 99, and fromthere out of outlet 23.

In accordance with one aspect of the present invention, the undulatingmembrane pump described herein avoids thrombus formation by placing allmoving parts directly within the primary flow path, thereby reducing therisk of flow stagnation. Specifically, the moving components depicted inFIG. 11, including magnetic ring assembly 76, suspension rings 79 and80, posts 81 and membrane assembly 82 all are located within deliverychannel 100 and flow channel 101. Flow stagnation may further be avoidedby eliminating secondary flow paths that may experience significantlyslower flow rates.

Turning now to FIGS. 13 and 14, a lower portion of implantable pump 20,including flanged portion 87, membrane assembly 82 and lower housingportion 23 is shown. Delivery channel 100 may be in fluid communicationwith membrane assembly 82 and flow channel 101 which is defined by abottom surface of flanged portion 87 and the interior surface of lowerhousing portion 25. Flanged portion 87 may comprise feature 102 thatextends downward as the bottom of flanged portion 87 moves radiallyinward. The interior surface of lower housing portion 25 may also slopeupward as it extends radially inward. The combination of the upwardslope of the interior surface of lower housing portion 25 and the bottomsurface of flanged portion 87 moving downward narrows flow channel 101as the channel moves radially inwards from delivery channel 100 tocircular aperture 99 of membrane 97, which is disposed about pump outlet23.

As explained above, membrane assembly 82 may be suspended by posts 81within flow channel 101 below the bottom surface of flanged portion 87and above the interior surface of lower housing portion 25. Membraneassembly 82 may be free to move up and down in the vertical directionwithin flow channel 101, which movement is constrained only bysuspension rings 79 and 80. Membrane assembly 82 may be constrained fromtwisting, tilting or moving in any direction in flow channel 101 otherthan up and down by rigid posts 81 and by the suspension rings.

Flow channel 101 is divided by membrane 97 into an upper flow channeland a lower flow channel by membrane 97. The geometry of membrane 97 maybe angled such that when membrane assembly 82 is at rest, the topsurface of membrane 97 is parallel to the bottom surface of flangedportion 87 and the bottom surface of membrane 97 is parallel to theopposing surface of lower housing portion 25. Alternatively, membrane 97may be sized and shaped such that when membrane assembly 82 is at rest,the upper and lower flow channels narrow as they move radially inwardfrom delivery channel 100 to circular aperture 99 in membrane 97.

Referring now also to FIG. 14, as rigid membrane ring 96 is caused byposts 81 to move up and down in flow channel 101, the outermost portionof membrane 97 nearest rigid membrane ring 96, moves up and down withrigid membrane ring 96. Membrane 97, being flexible and having elasticproperties, gradually translates the up and down movement of themembrane portion nearest rigid membrane ring 96 along membrane 97towards circular aperture 99. This movement across flexible membrane 97causes wavelike deformations in the membrane which may propagate inwardsfrom rigid membrane ring 96 towards aperture 99.

The waves formed in the undulating membrane may be manipulated bychanging the speed at which rigid membrane ring 96 moves up and down aswell as the distance rigid membrane ring 96 moves up and down. Asexplained above, the amplitude and frequency at which rigid membranering 96 moves up and down is determined by the amplitude and frequencyat which magnetic ring assembly 76 reciprocates over electromagnetassembly 91 Accordingly, the waves formed in the undulating membrane maybe adjusted by changing the frequency and amplitude at which magneticring assembly 76 is reciprocated.

When blood is introduced into flow channel 101 from delivery channel100, the undulations in membrane 97 cause blood to be propelled towardcircular aperture 99 and out of pump housing 27 via outlet 23. Thetransfer of energy from the membrane to the blood is directed radiallyinward along the length of the membrane towards aperture 99, and propelsthe blood along the flow channel towards outlet 23 along both sides ofmembrane 97.

For example, when rigid membrane ring 96 moves downward in unison withmagnetic ring assembly 76, the upper portion of flow channel 101 neardelivery channel 100 expands, causing blood from delivery channel 100 tofill the upper portion of the flow channel near the outer region ofmembrane 97. As rigid membrane ring 96 moves upward, the upper portionof flow channel 101 begins to narrow near rigid membrane ring 96,causing wave-like deformations to translate across the membrane. As thewave propagates across membrane 97, blood in the upper portion of flowchannel 101 is propelled towards circular aperture and ultimately out ofpump housing 27 through outlet 23. Simultaneously, as rigid membranering 96 moves upwards, the lower portion of flow channel 101 nearest theouter portion of membrane 97 begins to enlarge, allowing blood fromdelivery channel 100 to flow into this region. Subsequently, when rigidmembrane ring 96 is again thrust downwards, the region of lower portionof flow channel 101 nearest outer portion of membrane 97 begins tonarrow, causing wave-like deformations to translate across the membranethat propel blood towards outlet 23.

By manipulating the waves formed in the undulating membrane by changingthe frequency and amplitude at which magnetic ring assembly 76 moves upand down, the pressure gradient within flow channel 101 and ultimatelythe flow rate of the blood moving through flow channel 101 may beadjusted. Appropriately controlling the movement of magnetic ringassembly 76 permits oxygen-rich blood to be effectively and safelypumped from the left ventricle to the aorta and throughout the body asneeded.

In addition to merely pumping blood from the left ventricle to theaorta, implantable pump 20 of the present invention may be operated toclosely mimic physiologic pulsatility, without loss of pump efficiency.In the embodiment detailed above, pulsatility may be achieved nearlyinstantaneously by changing the frequency and amplitude at whichmagnetic ring assembly 76 moves, to create a desired flow output, or byceasing movement of the magnetic ring assembly for a period time tocreate a period of low or no flow output. Unlike typical rotary pumps,which require a certain period of time to attain a set number ofrotations per minute to achieve a desired fluid displacement andpulsatility, implantable pump 20 may achieve a desired flow outputnearly instantaneously and similarly may cease output nearlyinstantaneously due to the very low inertia generated by the smallmoving mass of the moving components of the pump assembly. The abilityto start and stop on-demand permits rapid changes in pressure and flow.Along with the frequency and amplitude, the duty cycle, defined by thepercentage of time membrane 97 is excited over a set period of time, maybe adjusted to achieve a desired flow output and pulsatility, withoutloss of pump efficiency. Even holding frequency and amplitude constant,flow rate may be altered by manipulating the duty cycle between 0 and100%.

In accordance with another aspect of the invention, controller 30 may beprogrammed by programmer 50 to operate at selected frequencies,amplitudes and duty cycles to achieve a wide range of physiologic flowrates and with physiologic pulsatilities. For example, programmer 50 maydirect controller 30 to operate implantable pump 20 at a givenfrequency, amplitude and/or duty cycle during a period of time when apatient is typically sleeping and may direct controller 30 to operateimplantable pump 20 at a different frequency, amplitude and or dutycycle during time periods when the patient is typically awake.Controller 30 or implantable pump also may include an accelerometer orposition indicator to determine whether the patient is supine orambulatory, the output of which may be used to move from one set of pumpoperating parameters to another. When the patient experiences certaindiscomfort or a physician determines that the parameters are notoptimized, physician may alter one or more of at least frequency,amplitude and duty cycle to achieve the desired functionality.Alternatively, controller 30 or mobile device 60 may be configured toalter one or more of frequency, amplitude and duty cycle to suit thepatient's needs.

Implantable pump 20 further may comprise one or more additional sensorsfor adjusting flow output and pulsatility according to the demand of thepatient. Sensors may be incorporated into implantable pump 20 oralternatively or in addition to may be implanted elsewhere in or on thepatient. The sensors preferably are in electrical communication withcontroller 30, and may monitor operational parameters that measure theperformance of implantable pump 20 or physiological sensors that measurephysiological parameters of the patients such as heart rate or bloodpressure. By using one or more physiological sensors, pulsatile flow maybe synchronized with a cardiac cycle of the patient by monitoring bloodpressure or muscle contractions, for example, and synchronizing the dutycycle according to the sensed output.

Controller 30 may compare physiological sensor measurements to currentimplantable pump output. If it is determined by analyzing sensormeasurements that demand exceeds current output, frequency, amplitudeand/or duty cycle may be automatically adjusted to meet current demand.Similarly, the controller may determine that current output exceedsdemand and thus alter output by changing frequency, amplitude and/orduty cycle. Alternatively, or in addition to, when it is determined thatdemand exceeds current output, an alarm may sound from controller 30.Similarly, operational measurements from operational sensors may becompared against predetermined thresholds and where measurements exceedpredetermined thresholds or a malfunction is detected, an alarm maysound from controller 30.

Implantable pump 20 is sized and shaped to produce physiological flowrates, pressure gradients and pulsatility at an operating point at whichmaximum efficiency is achieved. Specially, implantable pump 20 may besized and shaped to produce physiological flow rates ranging from 4 to 6liters per minute at pressure gradients lower than a threshold valueassociated with hemolysis. Also, to mimic a typical physiological pulseof 60 beats per minute, implantable pump 20 may pulse about once persecond. To achieve such pulsatility, a duty cycle of 50% may be utilizedwith an “on” period of 0.5 seconds and an “off” period of 0.5 seconds.For a given system, maximum efficiency at a specific operatingfrequency, amplitude and voltage may be achieved while producing a flowrate of 4 to 6 liters per minute at a duty cycle of 50% by manipulatingone or more of the shape and size of blood flow channels, elasticproperties of the suspension rings, mass of the moving parts, membranegeometries, and elastic properties and friction properties of themembrane. In this manner, implantable pump 20 may be designed to producedesirable physiological outputs while continuing to function at optimumoperating parameters.

By adjusting the duty cycle, implantable pump 20 may be configured togenerate a wide range of output flows at physiological pressuregradients. For example, for an exemplary LVAD system configured toproduce 4 to 6 liters per minute at a duty cycle of 50%, optimaloperating frequency may be 120 Hz. For this system, flow output may beincreased to 10 liters per minute or decreased to 4 liters per minute,for example, by changing only the duty cycle. As duty cycle andfrequency operate independent of one another, duty cycle may bemanipulated between 0 and 100% while leaving the frequency of 120 Hzunaffected.

The implantable pump system described herein, tuned to achievephysiological flow rates, pressure gradients and pulsatility, alsoavoids hemolysis and platelet activation by applying low to moderateshear forces on the blood, similar to those exerted by a healthy heart.The moving components are rigidly affixed to one another and do notincorporate any parts that would induce friction, such as mechanicalbearings or gears. In the embodiment detailed above, delivery channel100 may be sized and configured to also avoid friction between movingmagnetic ring assembly 76, suspension rings 79 and 80, posts 81 andlower housing portion 25 by sizing the channel such that clearances ofat least 0.5 mm are maintained between all moving components. Similarly,magnetic ring assembly 76, suspension rings 79 and 80, and posts 81 allmay be offset from stator assembly 72 by at least 0.5 mm to avoidfriction between the stator assembly and the moving parts.

Referring now to FIGS. 15A and 15B, an alternative exemplary embodimentof the pump assembly of the present invention is described. Implantablepump 20′ is constructed similar to implantable pump 20 described inFIGS. 7, 8, and 12, in which similar components are identified withlike-primed numbers. Implantable pump 20′ is distinct from implantablepump 20 in that membrane assembly 82′ includes skirt 115 coupled tomembrane 97′. Skirt illustratively includes first portion 115 a andsecond portion 115 b. First portion 115 a of skirt 115 extends upwardwithin delivery channel 100′ toward inlet 21′ in a first direction,e.g., parallel to the longitudinal axis of stator assembly 72′ and/or tothe central axis of pump housing 27′. Second portion 115 b of skirt 115curves toward outlet 23′ such that second portion 115 b is coupled tomembrane 97′ so that membrane 97′ is oriented in a second direction,e.g., perpendicular to first portion 115 a of skirt 115. For example,skirt 115 may have a J-shaped cross-section, such that first portion 115a forms a cylindrical-shaped ring about stator assembly 72′ and secondportion 115 b has a predetermined radius of curvature which allows bloodto flow smoothly from delivery channel 100′ across skirt 115 to theouter edge of membrane 97′ and into flow channel 101′, while reducingstagnation of blood flow. Skirt 115 breaks flow recirculation of bloodwithin delivery channel 100′ and improves hydraulic power generated fora given frequency while minimizing blood damage. In addition, theJ-shape of skirt 115 around stator assembly 72′ is significantly morestiff than a planar rigid membrane ring, thereby reducing flexing andfatigue, as well as drag as the blood moves across membrane 97′.

Skirt 115 exhibits rigid properties under typical forces experiencedduring the full range of operation of the present invention and may bemade of a biocompatible metal, e.g., titanium. Skirt 115 is preferablyimpermeable such that blood cannot flow through skirt 115. Postreception sites 98′ may be formed into skirt 115 to engage membraneassembly 82′ with posts 81′. Alternatively, posts 81′ may be attached toskirt 115 in any other way which directly translates the motion ofmagnetic ring assembly 76′ to skirt 115.

As magnetic ring assembly 76′ moves up and down, the movement is rigidlytranslated by posts 81′ to J-shape of skirt 115 of membrane assembly82′. Given the rigidity of the posts, when magnetic ring assembly 76′travels a certain distance upward or downward, membrane assembly 82′ maytravel the same distance. For example, when magnetic ring assembly 76′travels 2 mm from a position near first electromagnetic coil 77′ to aposition near second electromagnetic coil 78′, membrane assembly 82′ mayalso travel 2 mm in the same direction. Similarly, the frequency atwhich magnetic ring assembly 76′ traverses the space between the firstand second electromagnetic coils may be the same frequency at whichmembrane assembly 82′ travels the same distance.

Skirt 115 may be affixed to membrane 97′ and hold membrane 97′ intension. Membrane 97′ may be molded directly onto skirt 115 or may beaffixed to skirt 115 in any way that holds membrane 97′ uniformly intension along its circumference. For example, skirt 115 may be coatedwith the same material used to form membrane 97′ and the coating onskirt 115 may be integrally formed with membrane 97′.

Blood may enter implantable pump 20′ from the left ventricle throughinlet cannula 21′ and flow downward along the pump assembly intodelivery channel 100′. As the blood moves down tapered section 83′, itis directed through gap 74′ and into a vertical portion of deliverychannel 100′ in the area between pump housing 27′ and actuator assembly95′. As shown in FIG. 15A, skirt 115 divides delivery channel 100′ intoupper delivery channel 100 a and lower delivery channel 100 b such thatblood flow through delivery channel 100′ is divided into flow channel101 a via upper delivery channel 100 a and flow channel 101 b via lowerdelivery channel 100 b, wherein flow channels 101 a and 101 b areseparated by membrane 97′. As will be understood by one of ordinaryskill in the art, the volume of blood flow through each of deliverychannels 100 a and 100 b may depend on the diameter of first portion 115a of skirt 115. For example, the larger the diameter of first portion115 a of skirt 115, the larger the volume of delivery channel 100 a andthe smaller the volume of delivery channel 100 b. The ratio of thevolume of delivery channel 100 a to the volume of delivery channel 100 bmay be, for example, 1:1, 1:2, 1:3, 1:4, 2:1, 3:1, 4:1, etc., dependingon the amount of desired blood flow on each surface of membrane 97′.

By directing blood from inlet cannula 21′ across skirt 115 withindelivery channel 100′, blood flow is divided into delivery channel 100 aand 100 b and to flow channels 101 a and 101 b, respectively, such thatblood flows across the upper and lower surfaces of membrane 97′ ofmembrane assembly 82′. For example, as shown in FIG. 16A, blood flowthrough a pump having a planar rigid membrane ring spaced apart arelatively small distance from the pump housing may allow unrestrictedblood flow across the upper surface of the flexible membrane whilerestricting blood flow across the lower surface of the flexiblemembrane. Whereas, as depicted in FIG. 16B, blood flow through a pumphaving a J-shaped skirt may be distributed across both the upper andlower sides of the flexible membrane at a desired ratio.

Referring back to FIG. 15A, by actuating electromagnetic coils 77′ and78′, membrane 97′ may be undulated within flow channels 101 a and 101 bto induce wavelike formations in membrane 97′ that move from the edge ofmembrane 97′ towards circular aperture 99′. Accordingly, when blood isdelivered to membrane assembly 82′ from delivery channel 100′, it may bepropelled radially along both the upper and lower surfaces of membrane97′ towards circular aperture 99′, and from there out of outlet 23′. Thedistribution of blood flow across the upper and lower surfaces ofmembrane 97′ reduces recirculation of blood within delivery channel101′, and reduces repeated exposure of blood to high shear stress areas,which results in remarkably improved hydraulic performance ofimplantable pump 20′.

Referring now to FIG. 17, the relationship between the maximum hydraulicpower of the pump system and the height of the J-shaped skirt isdescribed. As the height of the vertical portion of the skirt increases,the maximum hydraulic power of the pump increases at a nonlinear rate.For example, as shown in FIG. 17, operation of a pump having a planarrigid membrane ring at 60 Hz results in a maximum of 0.15 W of hydraulicpower, at 90 Hz results in a maximum of 0.47 W of hydraulic power, andat 120 Hz results in a maximum of 1.42 W of hydraulic power. Operationof a pump having a skirt with an extension height of 2 mm, measured fromthe top surface of the membrane ring to the top of the J-shaped skirt,at 60 Hz results in a maximum of 0.16 W of hydraulic power, at 90 Hzresults in a maximum of 0.85 W of hydraulic power, and at 120 Hz resultsin a maximum of 1.54 W of hydraulic power. Operation of a pump having askirt with an extension height of 4 mm at 60 Hz results in a maximum of0.43 W of hydraulic power, at 90 Hz results in a maximum of 1.06 W ofhydraulic power, and at 120 Hz results in a maximum of 2.44 W ofhydraulic power. Operation of a pump having a skirt with an extensionheight of 10 mm at 60 Hz results in a maximum of 0.75 W of hydraulicpower, at 90 Hz results in a maximum of 1.89 W of hydraulic power, andat 120 Hz results in a maximum of 4.03 W of hydraulic power. Operationof a pump having a skirt with an extension height of 18 mm at 60 Hzresults in a maximum of 1.16 W of hydraulic power, at 90 Hz results in amaximum of 3.08 W of hydraulic power, and at 120 Hz results in a maximumof 9.13 W of hydraulic power. As such, height of skirt 115 is preferablyat least 2 mm, and more preferably at least 4 mm, at least 10 mm, and/orat least 18 mm. Accordingly, implantable pump 20′ may be operated at asignificantly lower frequency to achieve the same hydraulic output as apump having a planar rigid membrane ring operating at a higherfrequency, while reducing blood damage and increasing fatigue life ofmembrane 97′ and the springs.

Referring now to FIG. 18, an alternative exemplary embodiment of thepump assembly of the present invention having a J-shaped skirt isdescribed. Implantable pump 20″ is constructed similar to implantablepump 20′ described in FIG. 15A, in which similar components areidentified with like-double primed numbers. In addition, implantablepump 20″ includes skirt 115′ which is constructed similar to skirt 115of FIG. 15A. Implantable pump 20″ is distinct from implantable pump 20′in that inlet 21″ is coupled to inflow cannula 116, and outlet 23″ iscoupled to outflow cannula 117 such that outflow cannula 117 is disposedcoaxially within inflow cannula 116, as described in U.S. PatentPublication No. 2017/0290967 to Botterbusch, the entire contents ofwhich are incorporated herein by reference. Accordingly, duringoperation, blood flows into inlet 21″ via inflow cannula 116, throughdelivery channel 100″ into flow channel 101″ across membrane 97″, andexits through outlet cannula 117 via outlet 23″.

Referring now to FIG. 19, another alternative exemplary embodiment ofthe pump assembly of the present invention is described. Implantablepump 20′″ is constructed similar to implantable pump 20′ described inFIGS. 15A and 15B, in which similar components are identified withlike-double primed numbers and like-triple primed numbers. Implantablepump 20′″ is distinct from implantable pump 20′ in that implantable pump20′″ includes rigid ring 118 fixed about stator assembly 72′″. Ring 118extends longitudinally within delivery channel 100′″, parallel to thelongitudinal axis of stator assembly 72′″ such that ring 118 forms acylindrical-shaped ring about stator assembly 72′″.

In addition, membrane assembly 82′″ of implantable pump 20′″ includesskirt 119 coupled to membrane 97″. The upper portion of skirt 119 issubstantially parallel to ring 118, and the lower portion of skirt 119curves toward outlet 23′″ such that skirt 119 is coupled to membrane97″, perpendicular to ring 118. For example, skirt 119 may have aJ-shaped cross-section, having a predetermined radius of curvature whichallows blood to flow smoothly from delivery channels 100 a″ and 100 b″across skirt 119 to the outer edge of membrane 97′″ within flow channel101′″, while reducing stagnation of blood flow. Together, ring 118 andskirt 119 breaks flow recirculation of blood within delivery channel100′″ and improves hydraulic power generated for a given frequency whileminimizing blood damage. The distance between ring 118 and skirt 119 asskirt 119 reciprocates in response to the magnetic field generated bymagnetic ring assembly 76′″ as described in further detail below, isminimized to reduce leakage of blood between delivery channels 100 a″and 100 b″, and to reduce blood damage. In addition, the J-shape ofskirt 119 is significantly more stiff than a planar rigid membrane ring,thereby reducing flexing and fatigue, as well as drag as the blood movesacross membrane 97′″.

Skirt 119 is preferably impermeable such that blood cannot flow throughskirt 119, and exhibits rigid properties under typical forcesexperienced during the full range of operation of the present inventionand may be made of a biocompatible metal, e.g., titanium. Post receptionsites may be formed into skirt 119 to engage membrane assembly 82′″ withthe posts. Alternatively, the posts may be attached to skirt 119 in anyother way which directly translates the motion of magnetic ring assembly76′″ to skirt 119.

As magnetic ring assembly 76′″ moves up and down, the movement isrigidly translated by the posts to skirt 119 of membrane assembly 82′″.Given the rigidity of the posts, when magnetic ring assembly 76′″travels a certain distance upward or downward, membrane assembly 82′″may travel the same distance. For example, when magnetic ring assembly76′″ travels 2 mm from a position near first electromagnetic coil 77′″to a position near second electromagnetic coil 78″, membrane assembly82′″ may also travel 2 mm in the same direction. Similarly, thefrequency at which magnetic ring assembly 76′″ traverses the spacebetween the first and second electromagnetic coils may be the samefrequency at which membrane assembly 82′″ travels the same distance.

Skirt 119 may be affixed to membrane 97′″ and hold membrane 97′″ intension. Membrane 97′″ may be molded directly onto skirt 119 or may beaffixed to skirt 119 in any way that holds membrane 97′″ uniformly intension along its circumference. For example, skirt 119 may be coatedwith the same material used to form membrane 97′″ and the coating onskirt 119 may be integrally formed with membrane 97′″.

Blood may enter implantable pump 20′″ from the left ventricle throughinlet 21′″ and flow downward along the pump assembly into deliverychannel 100′″. As the blood moves down tapered section 83′″, it isdirected through gap 74′″ and into a vertical portion of deliverychannel 100′″ in the area between pump housing 27′″ and actuatorassembly 95′″. As shown in FIG. 19, ring 118 divides delivery channel100′″ into upper delivery channel 100 a″ and lower delivery channel 100b″ such that blood flow through delivery channel 100′″ is divided intoflow channel 101 a″ via upper delivery channel 100 a″ and flow channel101 b″ via lower delivery channel 100 b″ and across skirt 119 withminimal leakage between delivery channel 100 a″ and delivery channel 100b″, wherein flow channels 101 a″ and 101 b″ are separated by membrane97″_(.)

As will be understood by one of ordinary skill in the art, the volume ofblood flow through each of delivery channels 100 a″ and 100 b″ maydepend on the diameter of ring 118 and the curvature of radius of skirt119. For example, the larger the diameter of ring 118, the larger thevolume of delivery channel 100 a″ and the smaller the volume of deliverychannel 100 b″. The ratio of the volume of delivery channel 100 a″ tothe volume of delivery channel 100 b″ may be, for example, 1:1, 1:2,1:3, 1:4, 2:1, 3:1, 4:1, etc., depending on the amount of desired bloodflow on each surface of membrane 97′″. By directing blood from inletcannula 21″ across ring 118 within delivery channel 100′″, blood flow isdivided into delivery channels 100 a″ and 100 b″ and across skirt 119 toflow channels 101 a″ and 101 b″, respectively, such that blood flowsacross the upper and lower surfaces of membrane 97′″ of membraneassembly 82′″.

By actuating electromagnetic coils 77′″ and 78′″, membrane 97′″ may beundulated within flow channels 101 a″ and 101 b″ to induce wavelikeformations in membrane 97′″ that move from the edge of membrane 97′″towards circular aperture 99′″. Accordingly, when blood is delivered tomembrane assembly 82′″ from delivery channel 100′″, it may be propelledradially along both the upper and lower surfaces of membrane 97′″towards circular aperture 99′″, and from there out of outlet 23′″. Thedistribution of blood flow across the upper and lower surfaces ofmembrane 97′″ reduces recirculation of blood within delivery channel101′″, and reduces repeated exposure of blood to high shear stressareas, which results in remarkably improved hydraulic performance ofimplantable pump 20′″.

Referring now to FIG. 20, yet another alternative exemplary embodimentof the pump assembly of the present invention is described. Implantablepump 20″″ is constructed similar to implantable pump 20′″ described inFIG. 19, in which similar components are identified with like-primed,like-triple primed, and like-quadruple primed numbers. Implantable pump20″″ is distinct from implantable pump 20′″ in that implantable pump20″″ includes expandable portion 120 coupled between ring 118′ and theupper portion of skirt 119′. Expandable portion 120 is impermeable andprevents leakage between delivery channels 100 a′″ and 100 b′″.Preferably, expandable portion 120 has a pleated configuration that mayexpand and contract to permit efficient reciprocation of skirt 119′relative to ring 118′. For example, expandable portion 120 may comprisea plurality of bellows having a first end coupled to ring 118 ‘ and asecond end coupled to skirt 119’.

Expandable portion 120 may be molded directly onto skirt 119′ or may beaffixed to skirt 119′ in any way that holds expandable portion 120uniformly along its circumference. Similarly, expandable portion 120 maybe molded directly onto ring 118′ or may be affixed to ring 118′ in anyway that holds expandable portion 120 uniformly along its circumference.Skirt 119′ may be coated with the same material used to form membrane97′″ and/or expandable portion 120 and the coating on skirt 119′ may beintegrally formed with membrane 97′″ and/or expandable portion 120.

As shown in FIG. 20, expandable portion 120 extends longitudinallywithin delivery channel 100″″, parallel to the longitudinal axis ofstator assembly 72″″. Thus, during operation, blood is directed frominlet cannula 21″″ across ring 118′ and expandable portion 120 withindelivery channel 100″″, and divided into delivery channels 100 a′″ and100 b′″ and across skirt 119′ to flow channels 101 a′″ and 101 b′″,respectively, such that blood flows across the upper and lower surfacesof membrane 97″ of membrane assembly 82″.

As magnetic ring assembly 76″ moves up and down, the movement is rigidlytranslated by the posts to skirt 119′ of membrane assembly 82″, andthereby to expandable portion 120. For example, when magnetic ringassembly 76″ travels a certain distance upward or downward, membraneassembly 82″ travels the same distance causing expandable portion 120 toexpand and contract within delivery channel 100″″ parallel to thelongitudinal axis of stator assembly 72″ by the same distance.Similarly, the frequency at which magnetic ring assembly 76″ traversesthe space between the first and second electromagnetic coils may be thesame frequency at which membrane assembly 82″ travels the same distance.

Referring now to FIGS. 21A-21H, various configurations for energizingthe implantable pumps of the present invention, e.g., implantable pumps20, 20″, 20′″, and 20″″, described above are provided. As shown in FIG.21A, controller 30 includes output port 33 which is electrically coupledto cable 29 as described above, which in turn is coupled to theimplantable pump. Controller 30 also includes power connector 103, whichmay be electrically coupled to a battery, an extension port electricallycoupled to a battery, or an AC/DC power supply. For example, powerconnector 103 may be male, while the connector of the correspondingbattery or extension port is female.

In one embodiment, as shown in FIG. 21B, controller 30 includes twopower connectors, e.g., first power connector 103 and second powerconnector 104. As described above, first power connector 103 may beelectrically coupled to a first battery, a first extension portelectrically coupled to a first battery, or a first AC/DC power supply,and second power connector 103 may be electrically coupled to a secondbattery, a second extension port electrically coupled to a secondbattery, or a second AC/DC power supply. In this embodiment, first powerconnector 103 and second power connector 104 may both be male. Inaddition, controller 30 includes circuitry for switching between powersources such that energy is selectively transmitted to controller 30from at least one of the first or second battery/power supply. Forexample, the circuitry may switch between a first and second batteryintermittently, or after the remaining power level of one of thebatteries reaches a predetermined threshold.

Referring now to FIGS. 21C-E, configurations are illustrated whereincontroller 30 is directly electrically coupled to battery 40, such thatcontroller 30 and battery 40 may be worn by the patient together, e.g.,via a purse, shoulder bag, or holster. As shown in FIG. 21C, controller30 of FIG. 21A may be electrically coupled to battery 40 via powerconnector 103, wherein power connector 103 is male and battery 40 has acorresponding female connector. For example, FIG. 21D illustratescontroller 30 electrically coupled to battery 40, wherein battery 40 hasa smaller size, and therefore lower capacity, and FIG. 21E illustratescontroller 30 electrically coupled to battery 40, wherein battery 40 hasa larger size, and therefore higher capacity. As will be understood by aperson of ordinary skill in the art, battery 40 may have various sizesdepending on the need of the patient.

Referring now to FIGS. 21F-H, configurations are illustrated whereincontroller 30 is remotely electrically coupled to battery 40, such thatthe weight and volume of controller 30 and battery 40 are distributedand may be worn by the patient separately, e.g., via a belt or a vest.As shown in FIG. 21F, cable 41, which electrically couples controller 30to battery 40, is electrically coupled to first power connector port 105via strain relief 106, which is a hardwired junction between cable 41and first power connector port 105. Power connector port 105 includespower connector 107, which may be electrically coupled to a battery. Forexample, power connector 107 may be male, while the connector of thecorresponding battery is female.

As shown in FIG. 21G, controller 30 may be remotely electrically coupledto battery 40 via cable 41. Cable 41 is electrically coupled at one endto controller 30 via second power connector port 108 and strain relief114, which is a hardwired junction between cable 41 and second powerconnector port 108, and electrically coupled at another end to battery40 via first connector port 105 and strain relief 106. For example,power connector 103 of controller 30 may be male while the connector ofcorresponding second power connector port 108 is female, and powerconnector 107 of first power connector port 105 may be male while theconnector of corresponding battery 40 is female.

In one embodiment, as shown in FIG. 21H, controller 30 may be remotelyelectrically coupled to multiple batteries, e.g., battery 40A andbattery 40B, via a single second power connector port 108. As shown inFIG. 21H, second power connector port 108 includes first strain relief114A and second strain relief 114B, such that controller 30 is remotelyelectrically coupled to battery 40A via cable 41A and remotelyelectrically coupled to battery 40B via cable 41B. Specifically, cable41A is electrically coupled at one end to controller 30 via second powerconnector port 108 and first strain relief 114A, and electricallycoupled at another end to battery 40A via first connector port 105A andstrain relief 106A, and cable 41B is electrically coupled at one end tocontroller 30 via second power connector port 108 and second strainrelief 114B, and electrically coupled at another end to battery 40B viafirst connector port 105B and strain relief 106B. In this embodiment,controller 30 may include circuitry for switching between battery 40Aand battery 40B such that energy is selectively transmitted tocontroller 30 from at least one of battery 40A and battery 40B. Forexample, the circuitry may switch between battery 40A and battery 40Bintermittently, or after the remaining power level of one of thebatteries reaches a predetermined threshold. Alternatively, controller30 may receive energy from battery 40A and battery 40B simultaneously.

In another embodiment, as shown in FIG. 21I, controller 30 iselectrically coupled to AC/DC power supply 109, which may be pluggedinto an electrical outlet via AC plug 113, e.g., when the patient isresting bedside. Specifically, AC/DC power supply 109 is electricallycoupled to controller 30 via cable 41, such that cable 41 iselectrically coupled at one end to controller 30 via second powerconnector port 108 and strain relief 114, and electrically coupled atanother end to AC/DC power supply 109 via first power supply port 110.In addition, AC/DC power supply 109 is electrically coupled to plug 113via cable 112 and second power supply port 111.

Controller 30 may include an internal battery, such that the internalbattery powers controller 30 and the implantable pump during the timerequired for battery 40 to be replaced and/or recharged. Accordingly,controller 30 may include circuitry for switching between power sourcessuch that energy is transmitted to controller 30 from the internalbattery while battery 40 is disconnected from controller 30, and frombattery 40 when battery 40 is electrically coupled to controller 30. Inaddition, the circuitry may allow battery 40 to charge the internalbattery while also energizing the implantable pump until the internalbattery is recharged to a desired amount, at which point the circuitryallows battery 40 to solely energize the implantable pump. Similarly,when controller 40 is electrically coupled to AC/DC power supply 109,the circuitry may allow AC/DC power supply 109 to charge the internalbattery while also energizing the implantable pump until the internalbattery is recharged to a desired amount, at which point the circuitryallows AC/DC power supply 109 to solely energize the implantable pump.

In accordance with some aspects of the present invention, systems andmethods for controlling an implantable pump constructed in accordancewith the principles of the present invention, e.g., implantable pumps20, 20″, 20′″, and 20″″, without requiring position, velocity, oracceleration sensors are provided. Specifically, an exemplarycontroller, e.g., controller 30, for the implantable pump may only relyon the actuator's current measurement. The controller is robust topressure and flow changes inside the pump head, and allows fast changeof pump's operation point. For example, the controller includes, a twostage, nonlinear position observer module based on a reduced order modelof the electromagnetic actuator. As the actuator is very small regardingits performance requirements, linear approximation of the equivalentelectric circuit is insufficient. To meet the required operational rangeof the controller, the controller includes parameters' variationsregarding state variables. Means to identify the actuator's model aregiven by a recursive least squares (RLS) so they can be incorporated ina sensible way into the position observer module of the controller. Aforgetting factor is further included in the RLS to capture modelparameters' variations regarding state variables.

Referring now to FIG. 22, a flow chart illustrating steps of exemplarymethod 2200 for controlling an implantable pump constructed inaccordance with the principles of the present invention, e.g.,implantable pumps 20, 20″, 20′″, and 20″″. First, a finite elementmethod (FEM) model of an electromagnetic actuator, e.g., electromagnetassembly 91 and magnetic ring assembly 76, is transformed into a lumpedparameters model represented by a system of ordinary differentialequations (ODEs). The FEM model is set up by creating a subset of theactuator's geometry to save computing time as illustrated in FIG. 23. Atstep 2202, co-energy W of the implantable blood pump system is computedfor various magnetic ring positions and coil currents, as illustrated inFIG. 24A. For example, co-energy W may be approximated by a lookup tablethat stores the output co-energy W values of the FEM model simulation.

At step 2204, partial derivatives of co-energy W are computed andidentified to the parameters of an equivalent circuit which is expressedas:

${V_{in}(t)} = {{RI} + {{L( {x,I} )}\frac{dI}{dt}} + {{E( {x,I} )}\frac{dx}{dt}}}$where:${L( {x,I} )} = \frac{\partial^{2}{W( {x,I} )}}{\partial I^{2}}$${E( {x,I} )} = \frac{\partial^{2}W}{{\partial x}{\partial I}}$

The one degree-of-freedom motion equation of the magnetic ring of theimplantable pump gives:

$\begin{matrix}{{{m{\overset{¨}{x}(t)}} = {{F_{mag}( {x,I} )} + {F_{springs}(x)} + {F_{membrane}(t)}}}{{wher}\text{e:}}{F_{mag} = \frac{\partial{W( {x,I} )}}{\partial x}}{F_{springs} = {{ax^{3}} + {bx}}}} & \;\end{matrix}$

V_(in)=input voltagex=magnetic ring positionI=coil currentR=coil resistanceL=coil inductanceE=back EMF factor

FIGS. 24B, 24C, and 24D illustrate the force derived from the co-energyW as a function of the magnetic ring's position and coil current, theequivalent circuit's inductance derived from the co-energy W as afunction of the magnetic ring's position and coil current, and theequivalent circuit's EMF coefficient derived from the co-energy W as afunction of the magnetic ring's position and coil current, respectively.FIG. 25 is a graph illustrating the relationship between spring reactionforce and position of the magnetic ring of the implantable pump. FIG. 26is a schematic representation of the parameters of the equivalentcircuit and the one degree-of-freedom motion equation of the magneticring of the implantable pump described above.

Springs reaction force F_(springs) is identified to a third-degreepolynomial to take into account design-induced nonlinearities that aremeasured by the manufacturer of the electromagnetic actuator. Membraneforce F_(membrane) is supposed bounded and piecewise continuous. Thisvague description of the membrane force is motivated by the lack ofsufficient knowledge of the fluid-structure interaction that takes placebetween pump's membrane and fluid, as well as the possibility tosynthetize a controller that will not require more hypothesis of thisforce than what has been given.

At step 2206, the controller operates the electromagnetic actuator ofthe implantable pump to cause the moving component, e.g., magnetic ringassembly 76, to reciprocate at a predetermined stroke, e.g., frequencyand amplitude. At step 2208, the controller receives a signal indicativeof the alternating current of the system, e.g., coil current, from acurrent sensor positioned, for example, inside the power electronics ofthe implantable pump system.

For example, as illustrated in FIG. 27A, the implantable blood pump,which can be considered an inductive load, may be driven with an Hbridge configuration. As illustrated in schematic of the powerelectronics of FIG. 27A, the power electronics may include H bridge 130,shunt resistor 132, current sensor 140, and optional voltage sensor 150.The H bridge, illustrated in FIG. 27B, is driven to generate a certainvoltage waveform, while powered from a power supply, e.g., battery 40,directly or through a DC/DC voltage converter. As will be understood bya person having ordinary skill in the art, the power electronics couldinclude a single H bridge, with both coils in series or in parallel, ortwo H bridges, with one H bridge per coil.

FIG. 27C is a schematic of the current sensor for measuring the currentof the actuator. FIG. 27D is a schematic of an optional voltage sensorfor measuring the voltage of the actuator. The use of the voltagefeedback is optional, given that the algorithm controls the H bridge andtherefore knows the imposed voltage. As illustrated in FIG. 28, Analogto Digital Converter (ADC) sampling is synchronized with the middlepoint of the transistors pulse-width modulation (PWM) signals to removethe transistors switching glitch noise from both the current and voltagemeasurements. Accordingly, the current and/or pump voltage measurementsare sent to the algorithm running on the controller, and the algorithmestimates the position of the actuator and determines the H bridgevoltage required to impose a certain position oscillation.

Specifically, from the current measurement, the controller is able tocontrol the excitation of the deformable membrane, e.g., membrane 97,while being robust to the almost un-modelled force of the deformablemembrane F_(membrane). Thus, the implantable pump system may not requireposition, velocity, or acceleration sensors. For example, the controllerincludes a position observer module that has two stages. During thefirst stage (step 2210), the position observer module estimates thevelocity of the magnetic ring based on the alternating currentmeasurement and the parameters of an equivalent circuit using theequation described above:

${V_{in}(t)} = {{RI} + {{L( {x,I} )}\frac{dI}{dt}} + {{E( {x,I} )}\frac{dx}{dt}}}$

For example, the estimated velocity may be expressed as:

$\hat{\overset{.}{x}} = {\frac{1}{E( {x,I} )}( {V_{in} - {RI} - {{L( {x,I} )}\frac{dI}{dt}}} )}$

The derivative in of the above equation will make the estimationextremely sensitive to measurement noise if left as it is. To deal withthis estimation problem a derivate estimator is developed:

${(t)} = {{- \frac{6}{T^{3}}}{\int_{t - T}^{t}{( {T - {2\tau}} ){I( {t - \tau} )}d\tau}}}$

where T is the length of an integration window. This estimation isstraightforward to implement as a discrete finite impulse response (FIR)filter by using the trapezoidal method:

${( {kT_{s}} )} = {{- \frac{6}{( {NT_{s}} )^{3}}}{\sum\limits_{i = 0}^{N}{{w_{i}( {{NT_{s}} - {2iT_{s}}} )}{I( {{kT_{s}} - {iT_{s}}} )}d\tau}}}$

where N is an integer chosen so that T=NT_(s),

$w_{0} = {w_{N} = \frac{T_{s}}{2}}$

and w_(i)=T_(s), i=1, . . . , N−1.

Next, the second stage of the position observer module is implemented(step 2212), where the position observer module determines the velocityof the magnetic ring based on the estimated velocity calculated duringstep 2210. For example, it follows that, if {tilde over (x)} and {dotover ({tilde over (x)})} are the observed position and velocity, theposition observer module could be expressed as:

$\begin{bmatrix}\overset{\sim}{\overset{.}{x}} \\\overset{\sim}{\overset{¨}{x}}\end{bmatrix} = {{A\begin{bmatrix}\overset{\sim}{x} \\\overset{\sim}{\overset{.}{x}}\end{bmatrix}} + \begin{bmatrix}0 \\{F(t)}\end{bmatrix} + {\begin{bmatrix}k_{1} \\k_{2}\end{bmatrix}( {\hat{\overset{.}{x}} - \hat{\overset{.}{x}}} )}}$

where A is a constant square matrix regrouping the linear terms of theestimated velocity above and F(t) is the function regrouping thenonlinear elements, and k₁ and k₂, two gains to be chosen to guarantee:

${\lim\limits_{tarrow\infty}\begin{bmatrix}{x - \overset{\sim}{x}} \\{\overset{.}{x} - \overset{\sim}{\overset{.}{x}}}\end{bmatrix}} = 0$

At step 2214, the position observer module of the controller determinesthe position of the magnetic ring based on the determined velocityabove. Accordingly, from the observed position of the magnetic ring, thestroke controller will be able to set the excitation of the deformablemembrane via the electromagnetic actuator to a desired frequency andamplitude, while limiting overshoot. Thus, at step 2216, the controllercancels errors due to un-modeled dynamics of the implantable pump tolimit overshoot. For example, as illustrated in FIG. 29A, the controllermay include a feedforward module and a PI controller module. Thefeedforward module takes as input the desired position x_(d) at eachtime step to compute input voltage as:

$V_{in} = {{RI}_{d} + {L\frac{{dI}_{d}}{t}} + {E\frac{dx_{d}}{dt}}}$${F_{mag}( {x_{d},I_{d}} )} = {{m{\overset{¨}{x}}_{d}} - F_{springs} - {\alpha\frac{dx_{d}}{dt}}}$

where I_(d) can be computed as:

I _(d)=Φ(x _(d) ,F _(mag))

The reference signal x_(d) is generated as:

x_(d)(t) = S(t)sin (φ(t)) φ(t) = 2π f(t)${H(s)} = {\frac{k_{f}^{3}}{( {s + k_{f}} )^{3}} = {{\frac{S}{S_{d}}(s)} = {\frac{f}{f_{d}}(s)}}}$

where k_(f) is a positive, real number that guarantee the stability ofH(s).

Then, the remaining errors due to un-modeled dynamics are cancelled byPI controller module by adjusting the excitation signal. This could beimplemented using various methods. For example, its instantaneous valuecould be directly modified, or alternatively, another method is tomodify its amplitude, or both its amplitude and its instantaneous value,on different feedback loops as illustrated in FIG. 29B. If the amplitudemodification is used, one way to estimate it is to define an amplitudeestimator Ŝ(t) that is valid if x(t) is sufficiently close to a sinusfunction:

${\hat{S}(t)} = \sqrt{{\hat{x}(t)}^{2} + {\hat{x}( {t - \frac{1}{4f}} )}^{2}}$

At step 2218, the controller adjusts operation of the electromagneticactuator to cause the magnetic ring to reciprocate at an adjustedfrequency and/or amplitude, thereby causing the deformable membrane toproduce an adjusted predetermined blood flow across the implantablepump.

To capture the variations of inductance and back EMF coefficient withmagnetic ring's position and coil current, a recursive least squareestimator is used by the controller.

The parameters R, L and E described above are unknown and slowly timevarying. The variables V_(in), I and x are piecewise continuous andbounded, and all equal to zero at t=0. The problem is set by integrating(1) over t:

$I = {{\frac{1}{L}{\int V_{in}}} + {\frac{R}{L}{\int I}} + {\frac{E}{L}x}}$

which can be expressed as:

y = Ψ^(T)θ $\Psi^{T} = \lbrack \begin{matrix}{{\int V_{in}}\ } & {\int I} & x\end{matrix}\  \rbrack$ $\theta = \begin{bmatrix}\frac{1}{L} & \frac{R}{L} & \frac{E}{L}\end{bmatrix}$

For each sample n>0:

θ̂_(n) = θ̂_(n − 1) + K_(n)(y_(n) − ŷ_(n)) ŷ_(n) = Ψ_(n)^(T)θ̂_(n − 1)K_(n) = Ψ_(n)Q_(n)$Q_{n} = \frac{P_{n - 1}}{\lambda + {\Psi_{n}^{T}P_{n - 1}\Psi_{n}}}$$P_{n} = {\frac{1}{\lambda}( {P_{n - 1} - \frac{P_{n - 1}\Psi_{n}\Psi_{n}^{T}P_{n - 1}}{\lambda + {\Psi_{n}^{T}P_{n - 1}\Psi_{n}}}} )}$

where λ is a forgetting factor chosen so λ<1, P₀ is the initialcovariance matrix, and {circumflex over (θ)}₀ is the initial estimate ofthe parameters.

The resulting estimation data is then fit to polynomials of appropriatedegree, and stored into lookup tables that associate for each (x, I)combination the corresponding inductance and emf factor. The lookuptables are used in the velocity estimator:

${\hat{L}( {x,I} )},{{\hat{E}( {x,I} )} = {{\begin{bmatrix}1 & \ldots & x^{m}\end{bmatrix}\begin{bmatrix}a_{00} & \ldots & a_{0,n} \\\vdots & \ddots & \vdots \\a_{m,0} & \ldots & a_{m,n}\end{bmatrix}}\begin{bmatrix}1 \\\vdots \\I^{n}\end{bmatrix}}}$

Experimental Results

A numerical model of the implantable pump and the controller was builtunder Matlab/Simulink to test the implementation of the controller andmodel parameters' identification. The actuator model is compared tomeasurement and adjusted accordingly. The springs' reaction force ismeasured by using a pull tester, which is also used to measure themagnetic force of the actuator by applying an arbitrary constantelectric current on the electromagnetic coils of the actuator whilemeasuring force. The back EMF coefficient was derived from the forcemeasured at different electric currents and magnetic ring positions.Electric inductance and resistance may be estimated with a LRC meterwhen the magnetic ring's motion is blocked to cancel the effect of theback EMF. As LRC meters' input current is limited (<20 mA), inductancemay only be estimated in this limited area. In general, the performancesof the real actuator are reduced compared to the model (lowerinductance, magnetic force and EMF). This may be due to an imperfectmanufacturing process, e.g., the winding of the coils. The membraneforce may be emulated by a viscous friction term that is a sensiblefirst approximation:

F _(membrane)(t)=μ(t){dot over (x)}

With these verifications, the parameters' variations are identified andthe controller structure is tested. In particular, different positionobserver implementations are compared to show the interest of usingvarying electric parameters instead of linear approximations.

The results of the identification are shown in FIGS. 30A-30C. Toguarantee a quick convergence, two excitation signals are applied to theactuator. A voltage excitation that contains a high frequency (500 Hz)square wave voltage makes the inductance's voltage to never be close tozero, and a low frequency (0.1 Hz) sinus component for the resistance'svoltage reach every position. To ensure that back EMF is represented inthe response, an external sinus force is simulated at 50 Hz. To filterout high frequency variations as well as eventual noise, while capturingthe low frequency variations of the parameters, λ is set to 0.999 via atrial and error approach. The recursive least square identification wasrun with different initial conditions. Measurement errors (noises, bias,gain) were simulated to verify that their effects would not hinderconvergence and help to diagnostic future experimental issues. The RLSalgorithm filters out high frequency noises very easily, but gain errorslead to over or underestimation of parameters while bias and lowfrequency noise increase estimation error over time.

A discrete version of the controller is implemented on Simulink toemulate what would be done by compiling it on a hardware target. As thefrequency response of the derivative estimator described above dependson the length of the integration window and the sampling rate, and thesignals to derivate may have frequency up to 100 Hz. We set T_(s)=2×10⁻⁵s and N=6 (i.e. integration window of 1.2×10⁻⁴ s), which is a good tradebetween noise attenuation and performance. FIGS. 31A and 31B shows theresponse of the actuator from startup at t=0 s to a nominal constantoperation point, i.e. a constant amplitude and frequency. Current andmagnetic ring position are both reasonably sinusoidal, and after atransition period, the amplitude of position reaches the desiredamplitude, and the position observer module output keeps track of thevariation of position.

FIGS. 32A and 32B shows two cases of change of operation points: achange of frequency and a change of stroke, which are two ways toincrease or decrease blood flow through by the pump. As shown in FIGS.32A and 32B, overshoots appear during the change of stroke. If not keptbelow a safe level, the overshoots could create overstress that coulddamage the membrane and the springs. Overshoot may be avoided by makingthe desired stroke signal change smoother.

FIGS. 33A-33D is a comparison between 4 position observer modules whichare different through their first stage. In each case, inductance andback EMF are implemented as constant approximation and as functions. Asthe controller must maintain stroke over a wide range of strokes andfrequencies, and as different reaction forces of the membrane areunknown, combinations of those three parameters must be tested toevaluate the performance of the controller. To do so the variation ofthe membrane force is emulated according to flow and pressure inside thepump head by varying μ, and an error variable e is created that isevaluated over a range of strokes, frequency and μ:

${e( {S_{d},f_{d}} )} = {\max\limits_{\mu}( {{S_{d} - {{0.5}( {{\max x} - {\min\; x}} )}}} )}$

where max x & min x are computed from one period of oscillation. Thisformulation of e can be compared to a maximal admissible error ε: everyoperation point [S_(d), f_(d)] which presents e<ε can be reached safely(stroke will be maintained to deliver the required flow without the riskof damaging the device by an overshoot). With this performanceindicator, it is observed that taking into account the variations of theinductance and back EMF in the velocity estimator results in an increaseof the operation range of the stroke controller.

Referring now to FIG. 34, an exemplary multistage control of acontroller relying on position measurement constructed in accordancewith the principles of the present invention is provided. Specifically,magnetic ring position may be measured via one or more sensors and usedto control operation of the pump. For example, a sensor such as a halleffect sensor may be coupled to a component that remains stationaryrelative to the stator assembly, e.g., the stator assembly itself or thehousing of the pump. In addition, a small, permanent magnet may becoupled to any mobile component of the pump that moves relative to thestator assembly, e.g., the magnetic ring assembly or the rigid membranering. Accordingly, the intensity of the magnetic field generated by thepermanent magnet may be measured via the hall effect sensor duringoperation of the pump. For example, as the permanent magnet moves closethe hall effect sensor, the intensity of the magnetic field generated bythe permanent magnet and measured by the sensor increases.

The control scheme illustrated in FIG. 34 based on pump dynamics may bedeveloped with the magnetic ring position measurement and the magneticcoils' current described above. First, an appropriate trajectory may bebased on the inverse dynamics of the pump system.

The appropriate trajectory may be written as:

[x _(d) {dot over (x)} _(d) I _(d)]^(T)

The dynamics of the pump may be written as:

$\begin{bmatrix}\overset{¨}{x} \\\overset{.}{I}\end{bmatrix} = \begin{matrix}\frac{{F_{a}( {x,I} )} - {{\mu(t)}\overset{.}{x}}}{m + {m_{f}(t)}} \\{{\frac{1}{L( {x,I} )}V_{in}} - {\frac{R}{L( {x,I} )}I} - {\frac{E( {x,I} )}{L( {x,I} )}\overset{.}{x}}}\end{matrix}$

Where F_(a) is the magnetic force generated by the actuator plus thereaction force of the suspension springs. Given the desired strokeamplitude and frequency S_(d) and f_(d), it is possible to generate afeasible trajectory for the pump. Such a trajectory exists if thedesired position x_(d) may be derived twice with respect to time, andthose derivatives v_(d) and {umlaut over (x)}_(d) are continuous. As theforce map F_(a)(x, I) is a diffeomorphism, there exists a map φ_(a)(x,F_(d)) such as:

I _(d)=φ_(a)(x,F _(d))

F _(d)=(m+m _(f) _(k) ){umlaut over (x)} _(d)+{circumflex over (μ)}_(k)v _(d)

With

and {circumflex over (μ)}_(k), two estimates of m_(f) and μ may becomputed by the Kalman filter block. Moreover, the feedforward voltageV_(ff) that is required for the pump to follow the desired trajectorymay be computed for each k>0 (each number is related to time ast=kT_(s), T_(s) is the sampling period of the controller) as:

V _(ff) =RI _(d) +E(x _(d) ,I _(d))v _(d) +L(x _(d) ,I _(d))İ _(d)

As any model is stained with errors, a feedback voltage alone is notenough to follow the trajectory with enough accuracy. Thus, to completethe control scheme of FIG. 34, a feedback voltage is added, which may bewritten as:

V _(fb) _(k) =Ke _(k)

With K=[k₁ k₂ k₃ k₄] and the tracking error e_(k) such as e_(k)^(T)=[x_(d)−{circumflex over (x)}_(k) v_(d)−{circumflex over (v)}_(k)I_(d)−I_(k) z_(k)] and z_(k)=z_(k−1)+T_(s)(x_(d)−x_(k)). {circumflexover (x)}_(k) and {circumflex over (v)}_(k) are estimates of x and {dotover (x)} that are computed by the extended Kalman filter. A suitablechoice of matrix K may be selected, e.g., by using the Linear QuadraticIntegral method, applied to a linear approximation of pump dynamics.

The purpose of the Extended Kalman filter is to compute an accurateestimation of the variables x and {dot over (x)} and time varyingparameters μ(t) and m_(f)(t), given measurements of x and I that arecorrupted by noise.

Taking the dynamics of the pump that are discretized using Euler'smethod, the estimate variables and parameters may be written in a vectoras:

X _(k) ^(T)=[x _(k) v _(k) μ_(k) m _(f) _(k) ].

As the position x_(k) is measured, x_(k)=CX_(k) with C=[1 0 0 0]. Foreach k>0 we have:

${\hat{X}}_{{k + 1}|k} = {\begin{bmatrix}{\overset{\hat{}}{x}}_{k|k} \\{\overset{\hat{}}{v}}_{k|k} \\{\overset{\hat{}}{\mu}}_{k|k} \\

\end{bmatrix} + {T_{s}\begin{bmatrix}v_{k|k} \\\frac{( {{F_{actuator}( {x_{k},I_{k}} )} - {\mu_{k}v_{k}}} )}{m + m_{f_{k}}} \\0 \\0\end{bmatrix}}}$ P_(k + 1|k) = J_(k)P_(k|k)J_(k)^(T) + Q

where {circumflex over (X)}_(k+i|k) is a predicted estimate of vectorX_(k) ^(T) and P_(k+i|k) is the predicted covariance matrix. J_(k) isthe jacobian matrix:

$J_{k} = \begin{bmatrix}1 & {Ts} & 0 & 0 \\{\frac{T_{s}}{m + m_{f_{k❘k}}}\frac{\partial{F_{a}( {{\overset{\hat{}}{x}}_{k|k},I_{k}} )}}{\partial x}} & {{{- \frac{T_{s}}{m + m_{f_{k|k}}}}\mu_{k|k}} + 1} & {{- \frac{T_{s}}{m + m_{f_{k|k}}}}v_{k|k}} & {{- \frac{T_{s}}{( {m + m_{f_{k|k}}} )^{2}}}( {{F_{a}( {{\overset{\hat{}}{x}}_{k|k},I_{k}} )} - {\mu_{k|k}v_{k|k}}} )} \\0 & 0 & 1 & 0 \\0 & 0 & 0 & 1\end{bmatrix}$

And Q is a process covariance matrix made of 4 diagonal termsq_(1, . . . , 4)>0:

$Q = \begin{bmatrix}q_{1} & 0 & 0 & 0 \\0 & q_{2} & 0 & 0 \\0 & 0 & q_{3} & 0 \\0 & 0 & 0 & q_{4}\end{bmatrix}$

With the measurement covariance R>0 the correction gain L_(k+1), thecorrected estimate {circumflex over (X)}_(k+1|k+1) and covariance matrixP_(k+1|k+1) are computed as:

L_(k + 1) = P_(k + 1|k)C^(T)(CP_(k + 1|k)C^(T) + R)⁻¹X̂_(k + 1|k + 1) = X̂_(k + 1|k) + L_(k)(x_(k) − x̂_(k + 1|k))$P_{{k + 1}❘{k + 1}} = {( {\begin{bmatrix}1 & 0 & 0 & 0 \\0 & 1 & 0 & 0 \\0 & 0 & 1 & 0 \\0 & 0 & 0 & 1\end{bmatrix} - {L_{k}C}} )P_{{k + 1}|k}}$

In addition, blood flow from the inlet out through the outlet of thepump may be estimated based on the Kalman filter estimation of μdescribed above. For example, as it has been demonstrated experimentallythat there exists a strong correlation between the variation of{circumflex over (μ)} and the variation of pump flowrate q_(f) with agiven desired stroke and frequency, q_(f) may be written as a functionalsuch as:

q _(f) =f({circumflex over (μ)},S _(d) ,f _(d))

In accordance with this function, an estimation of pump flow may becomputed that does not require a flow sensor and which may be used toset the operating point of the pump.

Moreover, the Kalman filter estimation described above may be used todetect a fault of the pump system. Specifically, the estimation residualof the Kalman filter may be used to monitor the operating conditions ofthe pump at all times and may detect almost instantly any change thatwould necessarily be caused by a fault. Thus, for each k>0 the residualε_(k) may be computed as:

ε_(k) =x _(k) −{circumflex over (x)} _(k|k)

Thus, a statistical analysis of ε_(k) may be run to detect anomalies.For example, the average value of ε_(k) may be computed over anarbitrary number of samples. In nominal operation conditions, theaverage value of the residual should be close to zero. Accordingly, anytime the value is determined to be higher or lower than a thresholdvalue, a mechanical fault may be detected. In addition, abnormaloperation conditions may be detected by comparing the deviation of theestimated values and parameters with their expected nominal values.

Referring now to FIG. 35, another exemplary sensorless multistagecontrol of a controller constructed in accordance with the principles ofthe present invention is provided. The control scheme of FIG. 35 isconstructed similar to the control scheme of FIG. 34, except that theKalman filter of FIG. 35 does not rely on magnetic ring positionmeasurement. Instead, the Kalman filter of FIG. 35 relies on a velocityestimation

.

Taking the dynamics of the pump that are discretized using Euler'smethod, the estimate variables and parameters may be written as a vector

X _(k) ^(T)=[x _(k)

μ_(k) m _(f) _(k) ].

As the velocity

may be estimated by a velocity observer module,

=CX_(k) with C=[0 1 0 0]. For each k>0:

${\hat{X}}_{{k + 1}|k} = {\begin{bmatrix}{\overset{\hat{}}{x}}_{k|k} \\{\overset{\hat{}}{v}}_{k|k} \\{\overset{\hat{}}{\mu}}_{k|k} \\

\end{bmatrix} + {T_{s}\begin{bmatrix}v_{k|k} \\\frac{( {{F_{actuator}( {x_{k},I_{k}} )} - {\mu_{k}v_{k}}} )}{m + m_{f_{k}}} \\0 \\0\end{bmatrix}}}$ P_(k + 1|k) = J_(k)P_(k|k)J_(k)^(T) + Q

Where {circumflex over (X)}_(k+1|k) is a predicted estimate of vectorX_(k) ^(T) and P_(k+1|k) is the predicted covariance matrix. J_(k) isthe jacobian matrix:

$J_{k} = \begin{bmatrix}1 & {Ts} & 0 & 0 \\{\frac{T_{s}}{m + m_{f_{k❘k}}}\frac{\partial{F_{a}( {{\overset{\hat{}}{x}}_{k|k},I_{k}} )}}{\partial x}} & {{{- \frac{T_{s}}{m + m_{f_{k|k}}}}\mu_{k|k}} + 1} & {{- \frac{T_{s}}{m + m_{f_{k|k}}}}v_{k|k}} & {{- \frac{T_{s}}{( {m + m_{f_{k|k}}} )^{2}}}( {{F_{a}( {{\overset{\hat{}}{x}}_{k|k},I_{k}} )} - {\mu_{k|k}v_{k|k}}} )} \\0 & 0 & 1 & 0 \\0 & 0 & 0 & 1\end{bmatrix}$

And Q is a process covariance matrix made of 4 diagonal termsq_(1, . . . . , 4)>0:

$Q = \begin{bmatrix}q_{1} & 0 & 0 & 0 \\0 & q_{2} & 0 & 0 \\0 & 0 & q_{3} & 0 \\0 & 0 & 0 & q_{4}\end{bmatrix}$

With the measurement covariance R>0 the correction gain L_(k+1), thecorrected estimate {circumflex over (X)}_(k+1|k+1) and covariance matrixP_(k+1|k+1) may be computed as:

L_(k + 1) = P_(k + 1|k)C^(T)(CP_(k + 1|k)C^(T) + R)⁻¹X̂_(k + 1|k + 1) = X̂_(k + 1|k) + L_(k)(−v̂_(k + 1|k))$P_{{k + 1}❘{k + 1}} = {( {\begin{bmatrix}1 & 0 & 0 & 0 \\0 & 1 & 0 & 0 \\0 & 0 & 1 & 0 \\0 & 0 & 0 & 1\end{bmatrix} - {L_{k}C}} )P_{{k + 1}|k}}$

The velocity observer block may be computed in various ways. Forexample, the velocity observer block may be computed as:

$= \frac{V_{i\; n_{k}} - {{L( {{\overset{\hat{}}{x}}_{k},I_{k}} )}\frac{dI}{dt}} - {RI}_{k}}{E( {{\overset{\hat{}}{x}}_{k},I_{k}} )}$

$\frac{dI}{dt}$

may be approximated by the 1^(st) order finite difference method:

$\frac{dI}{dt} = \frac{( {I_{k} - I_{k - 1}} )}{T_{s}}$

Alternatively,

$\frac{dI}{dt}$

may be computed as:

${( {kT_{s}} )} = {{- \frac{6}{( {NT_{s}} )^{3}}}{\sum\limits_{i = 0}^{N}{{w_{i}( {{NT_{s}} - {2iT_{s}}} )}{I( {{kT_{s}} - {iT_{s}}} )}d\tau}}}$

N being an integer chosen so that T=NT_(s),

$w_{0} = {w_{N} = \frac{T_{s}}{2}}$

and w_(i)=T_(s), i=1, . . . N−1.

Alternatively, the velocity observer block may be computed using a setof nonlinear equations such as:

$\overset{\hat{}}{I} = {\int{( {{\frac{1}{L( {\overset{\hat{}}{x},I} )}V} - {\frac{R}{L( {\overset{\hat{}}{x},I} )}I} - \overset{\hat{}}{f}} )dt}}$$\overset{\hat{}}{f} = {{( {k_{1} + 1} )( {\overset{\hat{}}{I} - I + {\int{( {\overset{\hat{}}{I} - I} ){dt}}}} )} + {k_{2}{\int{{{sign}( {\overset{\hat{}}{I} - I} )}{dt}}}}}$$= {\frac{L( {\overset{\hat{}}{x},I} )}{E( {\overset{\hat{}}{x},I} )}\overset{\hat{}}{f}}$

Referring now to FIG. 36, yet another exemplary sensorless multistagecontrol of a controller constructed in accordance with the principles ofthe present invention is provided. As shown in FIG. 36, the estimatedposition of the position observer may be used in the extended Kalmanfilter of FIG. 34. Accordingly, the estimated velocity, position andparameters may be used in the control scheme of FIG. 36.

While various illustrative embodiments of the invention are describedabove, it will be apparent to one skilled in the art that variouschanges and modifications may be made therein without departing from theinvention. For example, pump assembly 70 shown in FIG. 9 may be ordereddifferently and may include additional or fewer components of varioussizes and composition. The appended claims are intended to cover allsuch changes and modifications that fall within the true spirit andscope of the invention.

What is claimed:
 1. An implantable blood pump system comprising: animplantable blood pump configured to be implanted at a patient's heart,the implantable blood pump comprising: a housing having an inlet and anoutlet; a deformable membrane disposed within the housing; and anactuator comprising a stationary component and a moving componentcoupled to the deformable membrane, the actuator configured to bepowered by an alternating current that causes the moving component toreciprocate at a predetermined frequency and amplitude relative to thestationary component, thereby causing the deformable membrane to producea predetermined blood flow from the inlet out through the outlet; and acontroller operatively coupled to the implantable blood pump, thecontroller configured to be programmed to: operate the actuator to causethe moving component to reciprocate at the predetermined frequency andamplitude relative to the stationary component; receive a signalindicative of the alternating current via a current sensor operativelycoupled to the controller; determine a position of the moving componentbased on the signal indicative of the alternating current; and adjustoperation of the actuator to cause the moving component to reciprocateat an adjusted predetermined frequency and amplitude relative to thestationary component based on the position of the moving component,thereby causing the deformable membrane to produce an adjustedpredetermined blood flow from the inlet out through the outlet.
 2. Thesystem of claim 1, wherein the controller is programmed to determine theposition of the moving component by estimating a velocity of the movingcomponent based on the signal indicative of the alternating current. 3.The system of claim 2, wherein the controller is programmed to estimatethe velocity of the moving component based on co-energy W values of afinite elements model (FEM) of various positions and alternatingcurrents of the moving component.
 4. The system of claim 2, wherein thecontroller is programmed to determine the position of the movingcomponent by determining the velocity of the moving component based onthe estimated velocity of the moving component.
 5. The system of claim1, wherein the controller is programmed to adjust operation of theactuator to cause the moving component to reciprocate at the adjustedpredetermined frequency and amplitude relative to the stationarycomponent while limiting overshoot.
 6. The system of claim 5, whereinthe controller comprises a PI controller configured to be programmed tolimit overshoot by canceling errors due to un-modeled dynamics of theimplantable blood pump.
 7. The system of claim 1, wherein the controlleris programmed to determine the position of the moving component based onthe signal indicative of the alternating current and variations ofinductance and back EMF coefficient.
 8. The system of claim 1, whereinthe adjusted predetermined blood flow comprises a pulse synchronizedwith the patient's heartbeat.
 9. The system of claim 1, wherein thestationary component comprises an electromagnetic assembly configured togenerate a magnetic field, and wherein the moving component comprises amagnetic ring concentrically suspended around the electromagneticassembly and configured to reciprocate responsive to the magnetic fieldat the predetermined frequency and amplitude over the electromagneticassembly.
 10. The system of claim 9, wherein the electromagneticassembly comprises first and second electromagnetic coils, and whereinthe magnetic ring is caused to move when at least one of the first orsecond electromagnetic coils is powered by the alternating current. 11.The system of claim 9, wherein the magnetic ring is configured to inducewave-like deformations in the deformable membrane by reciprocating overthe electromagnetic assembly.
 12. The system of claim 1, furthercomprising first and second suspension rings concentrically disposedaround and coupled to the stationary component and the moving component.13. The system of claim 12, wherein the moving component is coupled toeach of the deformable membrane and the first and second suspensionrings via a plurality of posts.
 14. The system of claim 12, wherein thefirst and second suspension rings permit the moving component toreciprocate relative to the stationary component.
 15. The system ofclaim 12, wherein the first and second suspension rings exert a springforce on the moving component when the moving component reciprocatesrelative to the stationary component.
 16. The system of claim 1, furthercomprising a rigid ring coupled to the moving component, the rigid ringfurther coupled to the deformable membrane.
 17. The system of claim 1,wherein a bottom surface of the actuator and an interior portion of thehousing adjacent the outlet forms a flow channel within which thedeformable membrane is suspended.
 18. The system of claim 17, whereinthe deformable membrane comprises a central aperture adjacent theoutlet.
 19. The system of claim 17, wherein the actuator and an interiorsurface of the housing adjacent the inlet forms a delivery channelextending from the inlet to the flow channel.
 20. The system of claim 1,further comprising a rechargeable battery configured to deliver thealternating current to power the implantable blood pump.